Time-of-flight measurement of skin or blood using array of laser diodes with bragg reflectors

ABSTRACT

A blood measurement system comprises an array of laser diodes, to generate light to penetrate tissue comprising skin, having one or more wavelengths, including a near-infrared wavelength, and Bragg reflector(s). At least one of the laser diodes to pulse at a pulse repetition rate between 1-100 megahertz. A detection system to measure blood in veins based at least in part on near-infrared diffuse reflection from the skin, the detection system comprising a photo-detector and a lens system coupled to the photo-detector, wherein the photo-detector is coupled to analog-to-digital converter(s) and a processor, and configured to measure absorption of hemoglobin in the near-infrared wavelength between 700-1300 nanometers, differentiate between regions in the skin with and without distinct veins, and implement pattern matching and a threshold function to correlate detected blood concentrations with a library of known concentrations to determine overlap.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.16/506,885 filed Jul. 9, 2019, now U.S. Pat. No. 10,517,484, which is acontinuation of U.S. application Ser. No. 16/272,069 filed Feb. 11,2019, which is a continuation of U.S. application Ser. No. 16/029,611filed Jul. 8, 2018, now U.S. Pat. No. 10,201,283, which is acontinuation of U.S. application Ser. No. 15/888,052 filed Feb. 4, 2018,now U.S. Pat. No. 10,136,819, which is a continuation of U.S.application Ser. No. 15/212,549 filed Jul. 18, 2016, now U.S. Pat. No.9,885,698, which is a continuation of U.S. application Ser. No.14/650,897 filed Jun. 10, 2015, now U.S. Pat. No. 9,494,567, which is aU.S. National Phase of PCT/US2013/075700 filed Dec. 17, 2013, whichclaims the benefit of U.S. provisional application Ser. No. 61/747,472filed Dec. 31, 2012, the disclosures of all of which are herebyincorporated in their entirety by reference herein.

U.S. application Ser. No. 16/506,885 is also a continuation of U.S.application Ser. No. 16/004,359 filed Jun. 9, 2018, which is acontinuation of U.S. application Ser. No. 14/109,007 filed Dec. 17, 2013(now U.S. Pat. No. 9,993,159), which claims the benefit of U.S.provisional application Ser. No. 61/747,553 filed Dec. 31, 2012, thedisclosures of all of which are hereby incorporated in their entirety byreference herein.

U.S. application Ser. No. 16/506,885 is also a continuation of U.S.application Ser. No. 16/188,194 filed Nov. 12, 2018, now U.S. Pat. No.10,386,230, which is a continuation of U.S. application Ser. No.16/004,154 filed Jun. 8, 2018 (now U.S. Pat. No. 10,126,283), which is acontinuation of U.S. application Ser. No. 15/855,201 filed Dec. 27, 2017(now U.S. Pat. No. 9,995,722), which is a continuation of U.S.application Ser. No. 15/711,907 filed Sep. 21, 2017 (now U.S. Pat. No.9,897,584), which is a divisional of U.S. application Ser. No.15/357,225 filed Nov. 21, 2016 (now U.S. Pat. No. 9,797,876), which is acontinuation of U.S. application Ser. No. 14/650,981 filed Jun. 10, 2015(now U.S. Pat. No. 9,500,634), which is the U.S. national phase of PCTApplication No. PCT/US2013/075767 filed Dec. 17, 2013, which claims thebenefit of U.S. provisional application Ser. No. 61/747,485 filed Dec.31, 2012, the disclosures of all of which are hereby incorporated byreference in their entirety.

U.S. application Ser. No. 16/506,885 is also a continuation of U.S.application Ser. No. 16/241,628 filed Jan. 7, 2019, now U.S. Pat. No.10,441,176, which is a continuation of U.S. Ser. No. 16/015,737 filedJun. 22, 2018 (now U.S. Pat. No. 10,172,523), which is a continuation ofU.S. Ser. No. 15/594,053 filed May 12, 2017 (now U.S. Pat. No.10,188,299), which is a continuation of U.S. application Ser. No.14/875,709 filed Oct. 6, 2015 (now U.S. Pat. No. 9,651,533), which is acontinuation of U.S. application Ser. No. 14/108,986 filed Dec. 17, 2013(now U.S. Pat. No. 9,164,032), which claims the benefit of U.S.provisional application Ser. No. 61/747,487 filed Dec. 31, 2012, thedisclosures of all of which are hereby incorporated in their entirety byreference herein.

U.S. application Ser. No. 16/506,885 is also a continuation of U.S.application Ser. No. 16/284,514 filed Feb. 25, 2019, which is acontinuation of U.S. application Ser. No. 16/016,649 filed Jun. 24, 2018(now U.S. Pat. No. 10,213,113), which is a continuation of U.S.application Ser. No. 15/860,065 filed Jan. 2, 2018 (now U.S. Pat. No.10,098,546), which is a Continuation of U.S. application Ser. No.15/686,198 filed Aug. 25, 2017 (now U.S. Pat. No. 9,861,286), which is acontinuation of U.S. application Ser. No. 15/357,136 filed Nov. 21, 2016(now U.S. Pat. No. 9,757,040), which is a continuation of U.S.application Ser. No. 14/651,367 filed Jun. 11, 2015 (now U.S. Pat. No.9,500,635), which is the U.S. national phase of PCT Application No.PCT/US2013/075736 filed Dec. 17, 2013, which claims the benefit of U.S.provisional application Ser. No. 61/747,477 filed Dec. 31, 2012 and U.S.provisional application Ser. No. 61/754,698 filed Jan. 21, 2013, thedisclosures of all of which are hereby incorporated by reference intheir entirety.

This application is related to U.S. provisional application Ser. No.61/747,477 filed Dec. 31, 2012; Ser. No. 61/747,481 filed Dec. 31, 2012;Ser. No. 61/747,485 filed Dec. 31, 2012; Ser. No. 61/747,487 filed Dec.31, 2012; Ser. No. 61/747,492 filed Dec. 31, 2012; Ser. No. 61/747,553filed Dec. 31, 2012; and Ser. No. 61/754,698 filed Jan. 21, 2013, thedisclosures of all of which are hereby incorporated in their entirety byreference herein.

This application is also related to International ApplicationPCT/US2013/075736 entitled Short-Wave Infrared Super-Continuum LasersFor Early Detection Of Dental Caries; U.S. application Ser. No.14/108,995 filed Dec. 17, 2013 entitled Focused Near-Infrared Lasers ForNon-Invasive Vasectomy And Other Thermal Coagulation Or OcclusionProcedures (U.S. Pat. App. Pub. No. US2014/0188092A1); InternationalApplication PCT/US2013/075767 entitled Short-Wave InfraredSuper-Continuum Lasers For Natural Gas Leak Detection, Exploration, AndOther Active Remote Sensing Applications; U.S. application Ser. No.14/108,986 filed Dec. 17, 2013 entitled Short-Wave InfraredSuper-Continuum Lasers For Detecting Counterfeit Or Illicit Drugs AndPharmaceutical Process Control (now U.S. Pat. No. 9,164,032); U.S.application Ser. No. 14/108,974 filed Dec. 17, 2013 entitledNon-Invasive Treatment Of Varicose Veins (U.S. Pat. App. Pub. No.US2014/018894A1); and U.S. application Ser. No. 14/109,007 filed Dec.17, 2013 entitled Near-Infrared Super-Continuum Lasers For EarlyDetection Of Breast And Other Cancers (now U.S. Pat. No. 9,993,159), thedisclosures of all of which are hereby incorporated in their entirety byreference herein.

BACKGROUND

With the growing obesity epidemic, the number of individuals withdiabetes is also increasing dramatically. For example, there are over200 million people who have diabetes. Diabetes control requiresmonitoring of the glucose level, and most glucose measuring systemsavailable commercially require drawing of blood. Depending on theseverity of the diabetes, a patient may have to draw blood and measureglucose four to six times a day. This may be extremely painful andinconvenient for many people. In addition, for some groups, such assoldiers in the battlefield, it may be dangerous to have to measureperiodically their glucose level with finger pricks.

Thus, there is an unmet need for non-invasive glucose monitoring (e.g.,monitoring glucose without drawing blood). The challenge has been that anon-invasive system requires adequate sensitivity and selectivity, alongwith repeatability of the results. Yet, this is a very large market,with an estimated annual market of over $10B in 2011 for self-monitoringof glucose levels.

One approach to non-invasive monitoring of blood constituents or bloodanalytes is to use near-infrared spectroscopy, such as absorptionspectroscopy or near-infrared diffuse reflection or transmissionspectroscopy. Some attempts have been made to use broadband lightsources, such as tungsten lamps, to perform the spectroscopy. However,several challenges have arisen in these efforts. First, many otherconstituents in the blood also have signatures in the near-infrared, sospectroscopy and pattern matching, often called spectral fingerprinting,is required to distinguish the glucose with sufficient confidence.Second, the non-invasive procedures have often transmitted or reflectedlight through the skin, but skin has many spectral artifacts in thenear-infrared that may mask the glucose signatures. Moreover, the skinmay have significant water and blood content. These difficulties becomeparticularly complicated when a weak light source is used, such as alamp. More light intensity can help to increase the signal levels, and,hence, the signal-to-noise ratio.

As described in this disclosure, by using brighter light sources, suchas fiber-based supercontinuum lasers, super-luminescent laser diodes,light-emitting diodes or a number of laser diodes, the near-infraredsignal level from blood constituents may be increased. By shining lightthrough the teeth, which have fewer spectral artifacts than skin in thenear-infrared, the blood constituents may be measured with lessinterfering artifacts. Also, by using pattern matching in spectralfingerprinting and various software techniques, the signatures fromdifferent constituents in the blood may be identified. Moreover,value-add services may be provided by wirelessly communicating themonitored data to a handheld device such as a smart phone, and thenwirelessly communicating the processed data to the cloud for storing,processing, and transmitting to several locations.

SUMMARY OF EXAMPLE EMBODIMENTS

In one embodiment, a smart phone or tablet comprises one or more laserdiodes configured to be pulsed and to generate light having one or moreoptical wavelengths, wherein at least a portion of the one or moreoptical wavelengths is a near-infrared wavelength between 700 nanometersand 2500 nanometers. A first one or more lenses is configured to receivea portion of the light from the one or more laser diodes and to directat least some portion of the received light to tissue. An array of laserdiodes is configured to be pulsed and to generate light having one ormore optical wavelengths, wherein at least a portion of the one or moreoptical wavelengths is a near-infrared wavelength between 700 nanometersand 2500 nanometers. A second one or more lenses is configured toreceive a portion of the light from the array of laser diodes, the arrayof laser diodes and the second one or more lenses configured to form thelight into a plurality of spots and to direct at least some of the spotsto tissue. An infrared camera is configured to be synchronized to the atleast one of the one or more laser diodes to receive at least a portionof light reflected from the tissue from at least one of the one or morelaser diodes, wherein the infrared camera generates data based at leastin part on the received light. The infrared camera is further configuredto be synchronized to the array of laser diodes to receive light from atleast a portion of the plurality of spots reflected from the tissue, andwherein the infrared camera generates additional data based at least inpart on the received light. The infrared camera is further configuredto: receive light while the one or more laser diodes and the array oflaser diodes are off and convert the received light into a first signal;and receive light while at least some of the one or more laser diodes orsome of the array of laser diodes are on, and convert the received lightinto a second signal, the received light including at least a part ofthe portion of the light from the at least one of the one or more laserdiodes reflected from the tissue, or at least a part of the portion ofthe light from the array of laser diodes reflected from the tissue. Thesmart phone or tablet is configured to generate a two-dimensional orthree-dimensional image using a difference between the first signal andthe second signal, and using at least part of the data or at least partof the additional data from the infrared camera. The smart phone ortablet further comprises a wireless receiver, a wireless transmitter, adisplay, a voice input module, and a speaker.

Embodiments may include a smart phone or tablet comprising one or morelaser diodes configured to be pulsed and to generate light having one ormore optical wavelengths, wherein at least a portion of the one or moreoptical wavelengths is a near-infrared wavelength between 700 nanometersand 2500 nanometers. A first one or more lenses is configured to receivea portion of the light from the one or more laser diodes and to directat least some portion of the received light to tissue. An array of laserdiodes is configured to be pulsed and to generate light having one ormore optical wavelengths, wherein at least a portion of the one or moreoptical wavelengths is a near-infrared wavelength between 700 nanometersand 2500 nanometers. A second one or more lenses is configured toreceive a portion of the light from the array of laser diodes, the arrayof laser diodes and the second one or more lenses configured to form thelight into a plurality of spots and to direct at least some of the spotsto tissue. An infrared camera is configured to be synchronized to the atleast one of the one or more laser diodes to receive at least a portionof light reflected from the tissue from at least one of the one or morelaser diodes, and wherein the infrared camera generates data based atleast in part on the received light. The infrared camera is furtherconfigured to be synchronized to the array of laser diodes to receivelight from at least a portion of the plurality of spots reflected fromthe tissue, wherein the infrared camera generates additional data basedat least in part on the received light. The smart phone or tablet isconfigured to generate a two-dimensional or three-dimensional imageusing at least part of the data or part of the additional data from theinfrared camera. The smart phone or tablet further comprises a wirelessreceiver, a wireless transmitter, a display, a voice input module, and aspeaker.

In one embodiment, a smart phone or tablet comprises one or more laserdiodes configured to be pulsed and to generate light having one or moreoptical wavelengths, wherein at least a portion of the one or moreoptical wavelengths is a near-infrared wavelength between 700 nanometersand 2500 nanometers. A first one or more lenses is configured to receivea portion of the light from the one or more laser diodes and to directat least some portion of the received light to tissue. An array of laserdiodes is configured to be pulsed and to generate light having one ormore optical wavelengths, wherein at least a portion of the one or moreoptical wavelengths is a near-infrared wavelength between 700 nanometersand 2500 nanometers. A second one or more lenses is configured toreceive a portion of the light from the array of laser diodes, the arrayof laser diodes and the second one or more lenses configured to form thelight into a plurality of spots and to direct at least some of the spotsto tissue, wherein the plurality of spots are also formed at least inpart by using an assembly in front of the array of laser diodes. Aninfrared camera is configured to be synchronized to the at least one ofthe one or more laser diodes to receive at least a portion of lightreflected from the tissue from at least one of the one or more laserdiodes, wherein the infrared camera generates data based at least inpart on the received light. The infrared camera is further configured tobe synchronized to the array of laser diodes to receive light from atleast a portion of the plurality of spots reflected from the tissue,wherein the infrared camera generates additional data based at least inpart on the received light. The smart phone or tablet is configured togenerate a two-dimensional or three-dimensional image using at leastpart of the data or part of the additional data from the infraredcamera. The smart phone or tablet further comprises a wireless receiver,a wireless transmitter, a display, a voice input module, and a speaker.

In one embodiment, an optical tomography system comprises an array oflaser diodes to generate light having one or more optical wavelengthsthat includes at least one near-infrared wavelength between 600nanometers and 1000 nanometers, at least one laser diode of the arraycomprises one or more Bragg reflectors. At least one of the laser diodesto pulse at a modulation frequency between 10 Megahertz and 1 Gigahertzand to have a phase associated with the modulation frequency, wherein atleast a portion of the light generated by the array is configured topenetrate tissue comprising skin. A detection system comprises at leastone photo-detector, a lens at an input to the at least onephoto-detector, and a processor to process digitized signals receivedfrom the at least one photo-detector, the detection system configured to(i) measure a phase shift of at least a portion of the light from thearray of laser diodes reflected from the tissue relative to the portionof the light generated by the array to penetrate the tissue, (ii)measure time-of-flight of at least a portion of the light from the arrayof laser diodes reflected from the tissue relative to the portion of thelight generated by the array to penetrate the tissue, (iii) generate oneor more images of the tissue based at least in part on an amplitude ofat least a portion of the light from the array of laser diodes reflectedfrom the tissue, and (iv) detect oxy- or deoxy-hemoglobin in the tissue.

In another embodiment, a blood measurement system comprises an array oflaser diodes to generate light having one or more optical wavelengthsthat includes at least one near-infrared wavelength, the array of laserdiodes comprising one or more Bragg reflectors, wherein at least aportion of the light generated by the array is configured to penetratetissue comprising skin. At least one of the laser diodes to pulse at apulse repetition rate between one (1) kilohertz and about 100 megahertz.A detection system to non-invasively measure blood in veins based atleast in part on near-infrared diffuse reflection from the skin, thedetection system comprising at least one photo-detector and a lenssystem coupled to the at least one photo-detector, wherein the at leastone photo-detector is coupled to one or more analog-to-digitalconverters and a processor. The detection system is configured tomeasure absorption of hemoglobin in the near-infrared wavelength between700 nanometers and 1300 nanometers. The detection system processor isconfigured to: differentiate between a region in the skin having a veinand a region in the skin without distinct veins, and implement patternmatching and a threshold function to correlate detected bloodconcentrations with a library of known concentrations to determineoverlap between the detected blood concentrations and knownconcentrations in the library.

In yet another embodiment an optical identification system comprises anarray of laser diodes to generate light having one or more opticalwavelengths that includes at least one near-infrared wavelength between600 nanometers and 1000 nanometers, at least one laser diode of thearray comprises one or more Bragg reflectors, wherein at least a portionof the light generated by the array of laser diodes is configured to bedirected to tissue comprising skin. At least one of the laser diodes topulse at a modulation frequency between 10 Megahertz and 1 Gigahertz andto have a phase associated with the modulation frequency. A detectionsystem comprises at least one photo-detector, a lens at an input to theat least one photo-detector, and a processor to process digitizedsignals received from the at least one photo-detector, the detectionsystem configured to (i) measure a phase shift of at least a portion ofthe light from the array of laser diodes reflected from the tissue, (ii)measure time-of-flight of at least a portion of the light from the arrayof laser diodes reflected from the tissue, (iii) generate one or moreimages of the tissue based at least in part on an amplitude of at leasta portion of the light from the array of laser diodes reflected from thetissue, and (iv) generate a depth image from a temporal distribution oflight relative to a reference signal.

BRIEF DESCRIPTION OF THE DRAWINGS

For a more complete understanding of the present disclosure, and forfurther features and advantages thereof, reference is now made to thefollowing description taken in conjunction with the accompanyingdrawings, in which:

FIG. 1 plots the transmittance versus wavenumber for glucose in themid-wave and long-wave infrared wavelengths between approximately 2.7 to12 microns.

FIG. 2 illustrates measurements of the absorbance of different bloodconstituents, such as glucose, hemoglobin, and hemoglobin A1c. Themeasurements are done using an FTIR spectrometer in samples with a 1 mmpath length.

FIG. 3A shows the normalized absorbance of water and glucose (not drawnto scale). Water shows transmission windows between about 1500-1850 nmand 2050-2500 nm.

FIG. 3B illustrates the absorbance of hemoglobin and oxygenatedhemoglobin overlapped with water.

FIG. 4A shows measured absorbance in different concentrations of glucosesolution over the wavelength range of about 2000 to 2400 nm. This datais collected using a SWIR super-continuum laser with the sample pathlength of about 1.1 mm.

FIG. 4B illustrates measured absorbance in different concentrations ofglucose solution over the wavelength range of about 1550 to 1800 nm. Thedata is collected using a SWIR super-continuum laser with a sample pathlength of about 10 mm.

FIG. 5 illustrates the spectrum for different blood constituents in thewavelength range of about 2 to 2.45 microns (2000 to 2450 nm).

FIG. 6 shows the transmittance versus wavelength in microns for theketone 3-hydroxybutyrate. The wavelength range is approximately 2 to 16microns.

FIG. 7 illustrates the optical absorbance for ketones as well as someother blood constituents in the wavelength range of about 2100 to 2400nm.

FIG. 8A shows the first derivative spectra of ketone and protein atconcentrations of 10 g/L (left). In addition, the first derivativespectra of urea, creatinine, and glucose are shown on the right atconcentrations of 10 g/L.

FIG. 8B illustrates the near infrared absorbance for triglyceride.

FIG. 8C shows the near-infrared reflectance spectrum for cholesterol.

FIG. 8D illustrates the near-infrared reflectance versus wavelength forvarious blood constituents, including cholesterol, glucose, albumin,uric acid, and urea.

FIG. 9 shows a schematic of the human skin. In particular, the dermismay comprise significant amounts of collagen, elastin, lipids, andwater.

FIG. 10 illustrates the absorption coefficients for water (includingscattering), adipose, collagen, and elastin.

FIG. 11 shows the dorsal of the hand, where a differential measurementmay be made to at least partially compensate for or subtract out theskin interference.

FIG. 12 shows the dorsal of the foot, where a differential measurementmay be made to at least partially compensate for or subtract out theskin interference.

FIG. 13 illustrates a typical human nail tissue structure and thecapillary vessels below it.

FIG. 14 shows the attenuation coefficient for seven nail samples thatare allowed to stand in an environment with a humidity level of 14%.These coefficients are measured using an FTIR spectrometer over thenear-infrared wavelength range of approximately 1 to 2.5 microns. Belowis also included the spectrum of glucose.

FIG. 15 illustrates the structure of a tooth.

FIG. 16A shows the attenuation coefficient for dental enamel and waterversus wavelength from approximately 600 nm to 2600 nm.

FIG. 16B illustrates the absorption spectrum of intact enamel anddentine in the wavelength range of approximately 1.2 to 2.4 microns.

FIG. 17 shows the near infrared spectral reflectance over the wavelengthrange of approximately 800 nm to 2500 nm from an occlusal tooth surface.The black diamonds correspond to the reflectance from a sound, intacttooth section. The asterisks correspond to a tooth section with anenamel lesion. The circles correspond to a tooth section with a dentinelesion.

FIG. 18A illustrates a clamp design of a human interface to cap over oneor more teeth and perform a non-invasive measurement of bloodconstituents.

FIG. 18B shows a mouth guard design of a human interface to perform anon-invasive measurement of blood constituents.

FIG. 19 illustrates a block diagram or building blocks for constructinghigh power laser diode assemblies.

FIG. 20 shows a platform architecture for different wavelength rangesfor an all-fiber-integrated, high powered, super-continuum light source.

FIG. 21 illustrates one embodiment of a short-wave infrared (SWIR)super-continuum (SC) light source.

FIG. 22 shows the output spectrum from the SWIR SC laser of FIG. 21 whenabout 10 m length of fiber for SC generation is used. This fiber is asingle-mode, non-dispersion shifted fiber that is optimized foroperation near 1550 nm.

FIG. 23 illustrates high power SWIR-SC lasers that may generate lightbetween approximately 1.4-1.8 microns (top) or approximately 2-2.5microns (bottom).

FIG. 24 schematically shows that the medical measurement device can bepart of a personal or body area network that communicates with anotherdevice (e.g., smart phone or tablet) that communicates with the cloud.The cloud may in turn communicate information with the user, healthcareproviders, or other designated recipients.

FIG. 25 shows the experimental set-up for a reflection-spectroscopybased stand-off detection system.

FIG. 26 shows what might be an eventual flow-chart of a smartmanufacturing process.

FIG. 27 illustrates the optical absorption of pure water, hemoglobinwithout oxygen, and hemoglobin saturated with oxygen.

FIG. 28 shows examples of various absorption bands of chemical speciesin the wavelength range between about 1200-2200 nm.

FIG. 29 depicts the structure of a female breast.

FIG. 30 illustrates particular embodiments of imaging systems foroptically scanning a breast.

FIG. 31 shows the normalized absorption spectra of main tissue absorbersin the NIR for breast cancer, between about 600-1100 nm.

FIG. 32 illustrates the normalized absorption coefficient in thewavelength range between about 500-1600 nm for many of the components ofbreast tissue.

FIGS. 33A and 33B show the typical spectra of the cancerous site of atreated rat and the corresponding site of a normal rat. FIG. 33Aillustrates logarithm of the inverse of reflection spectra; and FIG. 33Billustrates second derivative spectra.

FIG. 34 shows the second derivative of spectral changes over severalweeks between about 1600-1800 nm in rats with breast cancer.

FIG. 35 illustrates the second derivative spectra for cholesterol,collagen and elastin.

FIG. 36 shows the absorption coefficient as a function of wavelengthbetween about 1000 nm and 2600 nm for water, adipose and collagen.

FIG. 37 illustrates the absorbance for four types of collagen: collagenI, collagen II, collagen III, and collagen IV.

FIG. 38 shows an experimental set-up for testing chicken breast samplesusing collimated light. In this experiment, the collimated light has abeam diameter of about 3 mm.

FIG. 39 plots the measured depth of damage (in millimeters) versus thetime-averaged incident power (in Watts). Data is presented for laserwavelengths near 980 nm, 1210 nm and 1700 nm, and lines are drawncorresponding to penetration depths of approximately 2 mm, 3 mm, and 4mm.

FIG. 40 illustrates the optical absorption or density as a function ofwavelength between approximately 700 nm and 1300 nm for water,hemoglobin and oxygenated hemoglobin.

FIG. 41 shows a set-up used for in vitro damage experiments usingfocused infrared light. After a lens system, the tissue is placedbetween two microscope slides.

FIGS. 42A and 42B present histology of renal arteries comprisingendothelium, media and adventitia layers and some renal nerves in orbelow the adventitia.

FIG. 42A illustrates renal arteries with no laser exposure. FIG. 42Billustrates renal arteries after focused laser exposure, with the laserlight near 1708 nm.

FIG. 43 illustrates the experimental set-up for ex vivo skin lasertreatment with surface cooling to protect the epidermis and top layer ofthe dermis.

FIGS. 44A-44D show MTT histo-chemistry of ex vivo human skin treatedwith ˜1708 nm laser and cold window (5 seconds precool; 2 mm diameterspot exposure for 3 seconds) at 725 mW (FIGS. 44A and 44B) correspondingto ˜70 J/cm2 average fluence and 830 mW (FIGS. 44C and 44D)corresponding to ˜80 J/cm2 average fluence.

FIG. 45 exemplifies a dual-beam experimental set-up that may be used tosubtract out (or at least minimize the adverse effects of) light sourcefluctuations.

DETAILED DESCRIPTION

As required, detailed embodiments of the present disclosure aredisclosed herein; however, it is to be understood that the disclosedembodiments are merely exemplary of the disclosure that may be embodiedin various and alternative forms. The figures are not necessarily toscale; some features may be exaggerated or minimized to show details ofparticular components. Therefore, specific structural and functionaldetails disclosed herein are not to be interpreted as limiting, butmerely as a representative basis for teaching one skilled in the art tovariously employ the present disclosure.

Various ailments or diseases may require measurement of theconcentration of one or more blood constituents. For example, diabetesmay require measurement of the blood glucose and HbA1c levels. On theother hand, diseases or disorders characterized by impaired glucosemetabolism may require the measurement of ketone bodies in the blood.Examples of impaired glucose metabolism diseases include Alzheimer's,Parkinson's, Huntington's, and Lou Gehrig's or amyotrophic lateralsclerosis (ALS). Techniques related to near-infrared spectroscopy orhyper-spectral imaging may be particularly advantageous for non-invasivemonitoring of some of these blood constituents.

Hyper-spectral images may provide spectral information to identify anddistinguish between spectrally similar materials, providing the abilityto make proper distinctions among materials with only subtle signaturedifferences. In the SWIR wavelength range, numerous gases, liquids andsolids have unique chemical signatures, particularly materialscomprising hydro-carbon bonds, O—H bonds, N—H bonds, etc. Therefore,spectroscopy in the SWIR may be attractive for stand-off or remotesensing of materials based on their chemical signature, which maycomplement other imaging information.

One embodiment of remote sensing that is used to identify and classifyvarious materials is so-called “hyper-spectral imaging.” Hyper-spectralsensors may collect information as a set of images, where each imagerepresents a range of wavelengths over a spectral band. Hyper-spectralimaging may deal with imaging narrow spectral bands over anapproximately continuous spectral range. As an example, inhyper-spectral imaging the sun may be used as the illumination source,and the daytime illumination may comprise direct solar illumination aswell as scattered solar (skylight), which is caused by the presence ofthe atmosphere. However, the sun illumination changes with time of day,clouds or inclement weather may block the sun light, and the sun lightis not accessible in the night time. Therefore, it would be advantageousto have a broadband light source covering the SWIR that may be used inplace of the sun to identify or classify materials in remote sensing orstand-off detection applications.

In one embodiment, a SWIR camera or infrared camera system may be usedto capture the images. The camera may include one or more lenses on theinput, which may be adjustable. The focal plane assemblies may be madefrom mercury cadmium telluride material (HgCdTe), and the detectors mayalso include thermo-electric coolers. Alternately, the image sensors maybe made from indium gallium arsenide (InGaAs), and CMOS transistors maybe connected to each pixel of the InGaAs photodiode array. The cameramay interface wirelessly or with a cable (e.g., USB, Ethernet cable, orfiber optics cable) to a computer or tablet or smart phone, where theimages may be captured and processed. These are a few examples ofinfrared cameras, but other SWIR or infrared cameras may be used and areintended to be covered by this disclosure.

Described herein are just some examples of the beneficial use ofnear-infrared or SWIR lasers for active remote sensing or hyper-spectralimaging. However, many other spectroscopy and identification procedurescan use the near-infrared or SWIR light consistent with this disclosureand are intended to be covered by the disclosure. As one example, thefiber-based super-continuum lasers may have a pulsed output with pulsedurations of approximately 0.5-2 nsec and pulse repetition rates ofseveral Megahertz. Therefore, the active remote sensing orhyper-spectral imaging applications may also be combined with LIDAR-typeapplications. Namely, the distance or time axis can be added to theinformation based on time-of-flight measurements. For this type ofinformation to be used, the detection system would also have to betime-gated to be able to measure the time difference between the pulsessent and the pulses received. By calculating the round-trip time for thesignal, the distance of the object may be judged. In another embodiment,GPS (global positioning system) information may be added, so the activeremote sensing or hyper-spectral imagery would also have a location tagon the data. Moreover, the active remote sensing or hyper-spectralimaging information could also be combined with two-dimensional orthree-dimensional images to provide a physical picture as well as achemical composition identification of the materials. These are justsome modifications of the active remote sensing or hyper-spectralimaging system described in this disclosure, but other techniques mayalso be added or combinations of these techniques may be added, andthese are also intended to be covered by this disclosure.

Described herein are just some examples of the beneficial use ofnear-infrared or SWIR lasers for active remote sensing or hyper-spectralimaging. However, many other spectroscopy and identification procedurescan use the near-infrared or SWIR light consistent with this disclosureand are intended to be covered by the disclosure. As one example, thefiber-based super-continuum lasers may have a pulsed output with pulsedurations of approximately 0.5-2 nsec and pulse repetition rates ofseveral Megahertz. Therefore, the active remote sensing orhyper-spectral imaging applications may also be combined with LIDAR-typeapplications. Namely, the distance or time axis can be added to theinformation based on time-of-flight measurements. For this type ofinformation to be used, the detection system would also have to betime-gated to be able to measure the time difference between the pulsessent and the pulses received. By calculating the round-trip time for thesignal, the distance of the object may be judged. In another embodiment,GPS (global positioning system) information may be added, so the activeremote sensing or hyper-spectral imagery would also have a location tagon the data. Moreover, the active remote sensing or hyper-spectralimaging information could also be combined with two-dimensional orthree-dimensional images to provide a physical picture as well as achemical composition identification of the materials. These are justsome modifications of the active remote sensing or hyper-spectralimaging system described in this disclosure, but other techniques mayalso be added or combinations of these techniques may be added, andthese are also intended to be covered by this disclosure.

In some instances, it may be desirable to create multiple locations offocused light on the varicose vein. For example, the speed of thetreatment may be increased by causing thermal coagulation or occlusionat multiple locations. Multiple collimated or focused light beams may becreated in one assembly. In this embodiment, optionally a surfacecooling apparatus may be used, where a cooling fluid may be flowedeither touching or in close proximity to the skin. Also, in thisparticular embodiment a cylindrical assembly may optionally be used,where the cylindrical length may be several millimeters in length anddefined by a clamp or mount placed on or near the leg. In oneembodiment, a window and/or lenslet array is also shown on thecylindrical surface for permitting the light to be incident on the skinand varicose vein at multiple spots. The lenslet array may comprisecircular, spherical or cylindrical lenses, depending on the type ofspots desired. As before, one advantage of placing the lenslet array inclose proximity to the skin and varicose vein may be that a high NA,lens may be used. Also, the input from the lens and/or mirror assemblyto the lenslet array may be single large beam, or a plurality of smallerbeams. In one embodiment, a plurality of spots may be created by thelenslet array to cause a plurality of locations of thermal coagulationin the varicose vein. Any number of spots may be used and are intendedto be covered by this disclosure.

In a non-limiting example, a plurality of spots may be used, or whatmight be called a fractionated beam. The fractionated laser beam may beadded to the laser delivery assembly or delivery head in a number ofways. In one embodiment, a screen-like spatial filter may be placed inthe pathway of the beam to be delivered to the biological tissue. Thescreen-like spatial filter can have opaque regions to block the lightand holes or transparent regions, through which the laser beam may passto the tissue sample. The ratio of opaque to transparent regions may bevaried, depending on the application of the laser. In anotherembodiment, a lenslet array can be used at or near the output interfacewhere the light emerges. In yet another embodiment, at least a part ofthe delivery fiber from the infrared laser system to the delivery headmay be a bundle of fibers, which may comprise a plurality of fiber coressurrounded by cladding regions. The fiber cores can then correspond tothe exposed regions, and the cladding areas can approximate the opaqueareas not to be exposed to the laser light. As an example, a bundle offibers may be excited by at least a part of the laser system output, andthen the fiber bundle can be fused together and perhaps pulled down to adesired diameter to expose to the tissue sample near the delivery head.In yet another embodiment, a photonic crystal fiber may be used tocreate the fractionated laser beam. In one non-limiting example, thephotonic crystal fiber can be coupled to at least a part of the lasersystem output at one end, and the other end can be coupled to thedelivery head. In a further example, the fractionated laser beam may begenerated by a heavily multi-mode fiber, where the speckle pattern atthe output may create the high intensity and low intensity spatialpattern at the output. Although several exemplary techniques areprovided for creating a fractionated laser beam, other techniques thatcan be compatible with optical fibers are also intended to be includedby this disclosure.

Although the output from a fiber laser may be from a single ormulti-mode fiber, different spatial spot sizes or spatial profiles maybe beneficial for different applications. For example, in some instancesit may be desirable to have a series of spots or a fractionated beamwith a grid of spots. In one embodiment, a bundle of fibers or a lightpipe with a plurality of guiding cores may be used. In anotherembodiment, one or more fiber cores may be followed by a lenslet arrayto create a plurality of collimated or focused beams. In yet anotherembodiment, a delivery light pipe may be followed by a grid-likestructure to divide up the beam into a plurality of spots. These arespecific examples of beam shaping, and other apparatuses and methods mayalso be used and are consistent with this disclosure.

As used throughout this document, the term “couple” and or “coupled”refers to any direct or indirect communication between two or moreelements, whether or not those elements are physically connected to oneanother. As used throughout this disclosure, the term “spectroscopy”means that a tissue or sample is inspected by comparing differentfeatures, such as wavelength (or frequency), spatial location,transmission, absorption, reflectivity, scattering, refractive index, oropacity. In one embodiment, “spectroscopy” may mean that the wavelengthof the light source is varied, and the transmission, absorption orreflectivity of the tissue or sample is measured as a function ofwavelength. In another embodiment, “spectroscopy” may mean that thewavelength dependence of the transmission, absorption or reflectivity iscompared between different spatial locations on a tissue or sample. Asan illustration, the “spectroscopy” may be performed by varying thewavelength of the light source, or by using a broadband light source andanalyzing the signal using a spectrometer, wavemeter, or opticalspectrum analyzer.

As used throughout this document, the term “fiber laser” refers to alaser or oscillator that has as an output light or an optical beam,wherein at least a part of the laser comprises an optical fiber. Forinstance, the fiber in the “fiber laser” may comprise one of or acombination of a single mode fiber, a multi-mode fiber, a mid-infraredfiber, a photonic crystal fiber, a doped fiber, a gain fiber, or, moregenerally, an approximately cylindrically shaped waveguide orlight-pipe. In one embodiment, the gain fiber may be doped with rareearth material, such as ytterbium, erbium, and/or thulium. In anotherembodiment, the mid-infrared fiber may comprise one or a combination offluoride fiber, ZBLAN fiber, chalcogenide fiber, tellurite fiber, orgermanium doped fiber. In yet another embodiment, the single mode fibermay include standard single-mode fiber, dispersion shifted fiber,non-zero dispersion shifted fiber, high-nonlinearity fiber, and smallcore size fibers.

As used throughout this disclosure, the term “pump laser” refers to alaser or oscillator that has as an output light or an optical beam,wherein the output light or optical beam is coupled to a gain medium toexcite the gain medium, which in turn may amplify another input opticalsignal or beam. In one particular example, the gain medium may be adoped fiber, such as a fiber doped with ytterbium, erbium or thulium. Inone embodiment, the “pump laser” may be a fiber laser, a solid statelaser, a laser involving a nonlinear crystal, an optical parametricoscillator, a semiconductor laser, or a plurality of semiconductorlasers that may be multiplexed together. In another embodiment, the“pump laser” may be coupled to the gain medium by using a fiber coupler,a dichroic mirror, a multiplexer, a wavelength division multiplexer, agrating, or a fused fiber coupler.

As used throughout this document, the term “super-continuum” and or“supercontinuum” and or “SC” refers to a broadband light beam or outputthat comprises a plurality of wavelengths. In a particular example, theplurality of wavelengths may be adjacent to one-another, so that thespectrum of the light beam or output appears as a continuous band whenmeasured with a spectrometer. In one embodiment, the broadband lightbeam may have a bandwidth of at least 10 nm. In another embodiment, the“super-continuum” may be generated through nonlinear opticalinteractions in a medium, such as an optical fiber or nonlinear crystal.For example, the “super-continuum” may be generated through one or acombination of nonlinear activities such as four-wave mixing, the Ramaneffect, modulational instability, and self-phase modulation.

As used throughout this disclosure, the terms “optical light” and or“optical beam” and or “light beam” refer to photons or light transmittedto a particular location in space. The “optical light” and or “opticalbeam” and or “light beam” may be modulated or unmodulated. In oneembodiment, the “optical light” and or “optical beam” and or “lightbeam” may originate from a fiber, a fiber laser, a laser, a lightemitting diode, a lamp, a pump laser, or a light source. In general, the“near-infrared (NIR)” region of the electromagnetic spectrum coversbetween approximately 0.7 microns (700 nm) to about 2.5 microns (2500nm). However, it may also be advantageous to use just the short-waveinfrared between approximately 1.4 microns (1400 nm) and about 2.5microns (2500 nm). One reason for preferring the SWIR over the entireNIR may be to operate in the so-called “eye-safe” window, whichcorresponds to wavelengths longer than about 1400 nm. Therefore, for theremainder of the disclosure the SWIR will be used for illustrativepurposes. However, it should be clear that the discussion that followscould also apply to using the NIR wavelength range, or other wavelengthbands.

In this disclosure, the term “damage” refers to affecting a tissue orsample so as to render the tissue or sample inoperable. For instance, ifa particular tissue normally emits certain signaling chemicals, then by“damaging” the tissue is meant that the tissue reduces or no longeremits that certain signaling chemical. The term “damage” and or“damaged” may include ablation, melting, charring, killing, or simplyincapacitating the chemical emissions from the particular tissue orsample. In one embodiment, histology or histochemical analysis may beused to determine whether a tissue or sample has been damaged.

As used throughout this document, the terms “near” or “about” or thesymbol “˜” refer to one or more wavelengths of light with wavelengthsaround the stated wavelength to accomplish the function described. Forexample, “near 1720 nm” may include wavelengths of between about 1680 nmand 1760 nm. In one embodiment, the term “near 1720 nm” refers to one ormore wavelengths of light with a wavelength value anywhere betweenapproximately 1700 nm and 1740 nm. Similarly, as used throughout thisdocument, the term “near 1210 nm” refers to one or wavelengths of lightwith a wavelength value anywhere between approximately 1170 nm and 1250nm. In one embodiment, the term “near 1210 nm” refers to one or morewavelengths of light with a wavelength value anywhere betweenapproximately 1190 nm and 1230 nm.

Spectrum for Glucose

One molecule of interest is glucose. The glucose molecule has thechemical formula C₆H₁₂O₆, so it has a number of hydro-carbon bonds. Anexample of the infrared transmittance of glucose 100 is illustrated inFIG. 1. The vibrational spectroscopy shows that the strongest lines forbending and stretching modes of C—H and O—H bonds lie in the wavelengthrange of approximately 6-12 microns. However, light sources anddetectors are more difficult in the mid-wave infrared and long-waveinfrared, and there is also strongly increasing water absorption in thehuman body beyond about 2.5 microns. Although weaker, there are alsonon-linear combinations of stretching and bending modes between about 2to 2.5 microns, and first overtone of C—H stretching modes betweenapproximately 1.5-1.8 microns. These signatures may fall in valleys ofwater absorption, permitting non-invasive detection through the body. Inaddition, there are yet weaker features from the second overtones andhigher-order combinations between about 0.8-1.2 microns; in addition tobeing weaker, these features may also be masked by absorption in thehemoglobin. Hence, the short-wave infrared (SWIR) wavelength range ofapproximately 1.4 to 2.5 microns may be an attractive window fornear-infrared spectroscopy of blood constituents.

As an example, measurements of the optical absorbance 200 of hemoglobin,glucose and HbA1c have been performed using a Fourier-Transform InfraredSpectrometer—FTIR. As FIG. 2 shows, in the SWIR wavelength rangehemoglobin is nearly flat in spectrum 201 (the noise at the edges is dueto the weaker light signal in the measurements). On the other hand, theglucose absorbance 202 has at least five distinct peaks near 1587 nm,1750 nm, 2120 nm, 2270 nm and 2320 nm.

FIG. 3A overlaps 300 the normalized absorbance of glucose 301 with theabsorbance of water 302 (not drawn to scale). It may be seen that waterhas an absorbance feature between approximately 1850 nm and 2050 nm, butwater 302 also has a nice transmission window between approximately1500-1850 nm and 2050 to 2500 nm. For wavelengths less than about 1100nm, the absorption of hemoglobin 351 and oxygenated hemoglobin 352 inFIG. 3B has a number of features 350, which may make it more difficultto measure blood constituents. Also, beyond 2500 nm the water absorptionbecomes considerably stronger over a wide wavelength range. Therefore,an advantageous window for measuring glucose and other bloodconstituents may be in the SWIR between 1500 and 1850 nm and 2050 to2500 nm. These are exemplary wavelength ranges, and other ranges can beused that would still fall within the scope of this disclosure.

One further consideration in choosing the laser wavelength is known asthe “eye safe” window for wavelengths longer than about 1400 nm. Inparticular, wavelengths in the eye safe window may not transmit down tothe retina of the eye, and therefore, these wavelengths may be lesslikely to create permanent eye damage. The near-infrared wavelengthshave the potential to be dangerous, because the eye cannot see thewavelengths (as it can in the visible), yet they can penetrate and causedamage to the eye. Even if a practitioner is not looking directly at thelaser beam, the practitioner's eyes may receive stray light from areflection or scattering from some surface. Hence, it can always be agood practice to use eye protection when working around lasers. Sincewavelengths longer than about 1400 nm are substantially not transmittedto the retina or substantially absorbed in the retina, this wavelengthrange is known as the eye safe window. For wavelengths longer than 1400nm, in general only the cornea of the eye may receive or absorb thelight radiation.

Beyond measuring blood constituents such as glucose using FTIRspectrometers, measurements have also been conducted in anotherembodiment using super-continuum lasers, which will be described laterin this disclosure. In this particular embodiment, some of the exemplarypreliminary data for glucose absorbance are illustrated in FIGS. 4A and4B. The optical spectra 401 in FIG. 4A for different levels of glucoseconcentration in the wavelength range between 2000 and 2400 nm show thethree absorption peaks near 2120 nm (2.12 μm), 2270 nm (2.27 μm) and2320 nm (2.32 μm). Moreover, the optical spectra 402 in FIG. 4B fordifferent levels of glucose concentration in the wavelength rangebetween 1500 and 1800 nm show the two broader absorption peaks near 1587nm and 1750 nm. It should be appreciated that although data measuredwith FTIR spectrometers or super-continuum lasers have been illustrated,other light sources can also be used to obtain the data, such assuper-luminescent laser diodes, light emitting diodes, a plurality oflaser diodes, or even bright lamp sources that generate adequate lightin the SWIR.

Although glucose has a distinctive signature in the SWIR wavelengthrange, one problem of non-invasive glucose monitoring is that many otherblood constituents also have hydro-carbon bonds. Consequently, there canbe interfering signals from other constituents in the blood. As anexample, FIG. 5 illustrates the spectrum 500 for different bloodconstituents in the wavelength range of 2 to 2.45 microns. The glucoseabsorption spectrum 501 can be unique with its three peaks in thiswavelength range. However, other blood constituents such as triacetin502, ascorbate 503, lactate 504, alanine 505, urea 506, and BSA 507 alsohave spectral features in this wavelength range. To distinguish theglucose 501 from these overlapping spectra, it may be advantageous tohave information at multiple wavelengths. In addition, it may beadvantageous to use pattern matching algorithms and other software andmathematical methods to identify the blood constituents of interest. Inone embodiment, the spectrum may be correlated with a library of knownspectra to determine the overlap integrals, and a threshold function maybe used to quantify the concentration of different constituents. This isjust one way to perform the signal processing, and many othertechniques, algorithms, and software may be used and would fall withinthe scope of this disclosure.

Ketone Bodies Monitoring

Beyond glucose, there are many other blood constituents that may also beof interest for health or disease monitoring. In another embodiment, itmay be desirous to monitor the level of ketone bodies in the bloodstream. Ketone bodies are three water-soluble compounds that areproduced as by-products when fatty acids are broken down for energy inthe liver. Two of the three are used as a source of energy in the heartand brain, while the third is a waste product excreted from the body. Inparticular, the three endogenous ketone bodies are acetone, acetoaceticacid, and beta-hydroxybutyrate or 3-hydroxybutyrate, and the wasteproduct ketone body is acetone.

Ketone bodies may be used for energy, where they are transported fromthe liver to other tissues. The brain may utilize ketone bodies whensufficient glucose is not available for energy. For instance, this mayoccur during fasting, strenuous exercise, low carbohydrate, ketogenicdiet and in neonates. Unlike most other tissues that have additionalenergy sources such as fatty acids during periods of low blood glucose,the brain cannot break down fatty acids and relies instead on ketones.In one embodiment, these ketone bodies are detected.

Ketone bodies may also be used for reducing or eliminating symptoms ofdiseases or disorders characterized by impaired glucose metabolism. Forexample, diseases associated with reduced neuronal metabolism of glucoseinclude Parkinson's disease, Alzheimer's disease, amyotrophic lateralsclerosis (ALS, also called Lou Gehrig's disease), Huntington's diseaseand epilepsy. In one embodiment, monitoring of alternate sources ofketone bodies that may be administered orally as a dietary supplement orin a nutritional composition to counteract some of the glucosemetabolism impairments is performed. However, if ketone bodiessupplements are provided, there is also a need to monitor the ketonelevel in the blood stream. For instance, if elevated levels of ketonebodies are present in the body, this may lead to ketosis; hyperketonemiais also an elevated level of ketone bodies in the blood. In addition,both acetoacetic acid and beta-hydroxybutyric acid are acidic, and, iflevels of these ketone bodies are too high, the pH of the blood maydrop, resulting in ketoacidosis.

The general formula for ketones is C_(n)H_(2n0). In organic chemistry, aketone is an organic compound with the structure RC(══O)R′, where R andR′ can be a variety of carbon-containing substituents. It features acarbonyl group (C══O) bonded to two other carbon atoms. Because theketones contain the hydrocarbon bonds, there might be expected to befeatures in the SWIR, similar in structure to those found for glucose.

The infrared spectrum 600 for the ketone 3-hydroxybutyrate isillustrated in FIG. 6. Just as in glucose, there are significantfeatures in the mid- and long-wave infrared between 6 to 12 microns, butthese may be difficult to observe non-invasively. On the other hand,there are some features in the SWIR that may be weaker, but they couldpotentially be observed non-invasively, perhaps through blood and water.

The optical spectra 700 for ketones as well as some other bloodconstituents are exemplified in FIG. 7 in the wavelength range of 2100nm to 2400 nm. In this embodiment, the absorbance for ketones is 701,while the absorbance for glucose is 702. However, there are alsofeatures in this wavelength range for other blood constituents, such asurea 703, albumin or blood protein 704, creatinine 705, and nitrite 706.In this wavelength range of 2100 to 2400 nm, the features for ketone 701seem more spectrally pronounced than even glucose.

Different signal processing techniques can be used to enhance thespectral differences between different constituents. In one embodiment,the first or second derivatives of the spectra may enable betterdiscrimination between substances. The first derivative may help removeany flat offset or background, while the second derivative may help toremove any sloped offset or background. In some instances, the first orsecond derivative may be applied after curve fitting or smoothing thereflectance, transmittance, or absorbance. For example, FIG. 8Aillustrates the derivative spectra for ketone 801 and glucose 802, whichcan be distinguished from the derivative spectra for protein 803, urea804 and creatinine 805. Based on FIG. 8A, it appears that ketones 801may have a more pronounced difference than even glucose 802 in thewavelength range between 2100 and 2400 nm. Therefore, ketone bodiesshould also be capable of being monitored using a non-invasive opticaltechnique in the SWIR, and a different pattern matching library could beused for glucose and ketones.

Hemoglobin A1c Monitoring

Another blood constituent that may be of interest for monitoring ofhealth or diseases is hemoglobin A1c, also known as HbA1c or glycatedhemoglobin (glycol-hemoglobin or glycosylated hemoglobin). HbA1c is aform of hemoglobin that is measured primarily to identify the averageplasma glucose concentration over prolonged periods of time. Thus, HbA1cmay serve as a marker for average blood glucose levels over the previousmonths prior to the measurements.

In one embodiment, when a physician suspects that a patient may bediabetic, the measurement of HbA1c may be one of the first tests thatare conducted. An HbA1c level less than approximately 6% may beconsidered normal. On the other hand, an HbA1c level greater thanapproximately 6.5% may be considered to be diabetic. In diabetesmellitus, higher amounts of HbA1c indicate poorer control of bloodglucose levels. Thus, monitoring the HbA1c in diabetic patients mayimprove treatment. Current techniques for measuring HbA1c requiredrawing blood, which may be inconvenient and painful. The point-of-caredevices use immunoassay or boronate affinity chromatography, as anexample. Thus, there is also an unmet need for non-invasive monitoringof HbA1c.

FIG. 2 illustrates the FTIR measurements of HbA1c absorbance 203 overthe wavelength range between 1500 and 2400 nm for a concentration ofapproximately 1 mg/ml. Whereas the absorbance of hemoglobin 201 overthis wavelength range is approximately flat, the HbA1c absorbance 203shows broad features and distinct curvature. Although the HbA1cabsorbance 203 does not appear to exhibit as pronounced features asglucose 202, the non-invasive SWIR measurement should be able to detectHbA1c with appropriate pattern matching algorithms. Moreover, thespectrum for HbA1c may be further enhanced by using first or secondderivative data, as seen for ketones in FIG. 8A. Beyond absorption,reflectance, or transmission spectroscopy, it may also be possible todetect blood constituents such as HbA1c using Raman spectroscopy orsurface-enhanced Raman spectroscopy. In general, Raman spectroscopy mayrequire higher optical power levels.

As an illustration, non-invasive measurement of blood constituents suchas glucose, ketone bodies, and HbA1c has been discussed thus far.However, other blood constituents can also be measured using similartechniques, and these are also intended to be covered by thisdisclosure. In other embodiments, blood constituents such as proteins,albumin, urea, creatinine or nitrites could also be measured. Forinstance, the same type of SWIR optical techniques might be used, butthe pattern matching algorithms and software could use different libraryfeatures or functions for the different constituents.

In yet another embodiment, the optical techniques described in thisdisclosure could also be used to measure levels of triglycerides.Triglycerides are bundles of fats that may be found in the blood stream,particularly after ingesting meals. The body manufactures triglyceridesfrom carbohydrates and fatty foods that are eaten. In other words,triglycerides are the body's storage form of fat. Triglycerides arecomprised of three fatty acids attached to a glycerol molecule, andmeasuring the level of triglycerides may be important for diabetics. Thetriglyceride levels or concentrations in blood may be rated as follows:desirable or normal may be less than 150 mg/dl; borderline high may be150-199 mg/dl; high may be 200-499 mg/dl; and very high may be 500 mg/dlor greater. FIG. 8B illustrates one example of the near-infraredabsorbance 825 for triglycerides. There are distinct absorbance peaks inthe spectrum that should be measurable. The characteristic absorptionbands may be assigned as follows: (a) the first overtones of C—Hstretching vibrations (1600-1900 nm); (b) the region of second overtonesof C—H stretching vibrations (1100-1250 nm); and, (c) two regions(2000-2200 nm and 1350-1500 nm) that comprise bands due to combinationsof C—H stretching vibrations and other vibrational modes.

A further example of blood compositions that can be detected or measuredusing near-infrared light includes cholesterol monitoring. For example,FIG. 8C shows the near-infrared reflectance spectrum for cholesterol 850with wavelength in microns (μm). Distinct absorption peaks areobservable near 1210 nm (1.21 μm), 1720 nm (1.72 μm), and between2300-2500 nm (2.3-2.5 μm). Also, there are other features near 1450 nm(1.45 μm) and 2050 nm (2.05 μm). In FIG. 8D the near-infraredreflectances 875 are displayed versus wavelength (nm) for various bloodconstituents. The spectrum for cholesterol 876 is overlaid with glucose877, albumin 878, uric acid 879, and urea 880. As may be noted from FIG.8D, at about 1720 nm and 2300 nm, cholesterol 876 reaches approximatereflectance peaks, while some of the other analytes are in a moregradual mode. Various signal processing methods may be used to identifyand quantify the concentration of cholesterol 876 and/or glucose 877, orsome of the other blood constituents.

As illustrated by FIGS. 5 and 7, one of the issues in measuring aparticular blood constituent is the interfering and overlapping signalfrom other blood constituents. The selection of the constituent ofinterest may be improved using a number of techniques. For example, ahigher light level or intensity may improve the signal-to-noise ratiofor the measurement. Second, mathematical modeling and signal processingmethodologies may help to reduce the interference, such as multivariatetechniques, multiple linear regression, and factor-based algorithms, forexample. For instance, a number of mathematical approaches includemultiple linear regression, partial least squares, and principalcomponent regression (PCR). Also, as illustrated in FIG. 8A, variousmathematical derivatives, including the first and second derivatives,may help to accentuate differences between spectra. In addition, byusing a wider wavelength range and using more sampling wavelengths mayimprove the ability to discriminate one signal from another. Moreover,it may be advantageous to pulse the light source with a particular pulsewidth and pulse repetition rate, and then the detection system canmeasure the pulsed light returned from or transmitted through thetissue. Using a lock-in type technique (e.g., detecting at the samefrequency as the pulsed light source and also possibly phase locked tothe same signal), the detection system may be able to reject backgroundor spurious signals and increase the signal-to-noise ratio of themeasurement. In one particular embodiment, high signal-to-noise ratiomay be achieved using a fiber-based super-continuum (SC) light source(described further herein). Other light sources may also be used,including a plurality of laser diodes, super-luminescent laser diodes,or fiber lasers. In one embodiment, an all-fiber-integrated,high-powered SC light source may be elegant for its simplicity. Thelight may be first generated from a seed laser diode (LD). For example,the seed LD may be a distributed feedback (DFB) laser diode with awavelength near 1542 nm or 1550 nm, with approximately 0.5-2.0 ns pulsedoutput, and with a pulse repetition rate between one kilohertz and about100 MHz or more.

Beyond having a pulse width, the laser output can also have a preferredrepetition rate. For pulse repetition rates above around 10 MHz, wheremultiple pulses fall within a thermal diffusion time, the tissueresponse may be more related to the energy deposited or the fluence ofthe laser beam. The separation between pulses or a sub-group of pulsesmay also be selected so that the tissue sample can reach thermalequilibrium between pulses. Also, the pulse pattern may or may not beperiodic. In one embodiment, there may be several pulses used per spot,where the pulse pattern is selected to obtain a desired thermal profile.The laser beam may then be moved to a new spot and then another pulsetrain delivered to that spot. In one embodiment, there can be severalseconds of pre-cooling, the laser can be exposed on the tissue forseveral seconds, and then there may also be post-cooling. Althoughparticular examples of laser duration and repetition rate are described,other values may also be used consistent with this disclosure. Forexample, depending on the application and mechanisms, the pulse ratecould range all the way from continuous wave to 100's of Megahertz.

The above are just examples of some of the methods of improving theability to discriminate between different constituents, but othertechniques may also be used and are intended to be covered by thisdisclosure.

In one embodiment, a SWIR camera or infrared camera system may be usedto capture the images. The camera may include one or more lenses on theinput, which may be adjustable. The focal plane assemblies may be madefrom mercury cadmium telluride material (HgCdTe), and the detectors mayalso include thermo-electric coolers. Alternately, the image sensors maybe made from indium gallium arsenide (InGaAs), and CMOS transistors maybe connected to each pixel of the InGaAs photodiode array. The cameramay interface wirelessly or with a cable (e.g., USB, Ethernet cable, orfiber optics cable) to a computer or tablet or smart phone, where theimages may be captured and processed. These are a few examples ofinfrared cameras, but other SWIR or infrared cameras may be used and areintended to be covered by this disclosure.

By use of an active illuminator, a number of advantages may be achieved.First, the variations due to sunlight and time-of-day may be factoredout. The effects of the weather, such as clouds and rain, might also bereduced. Also, higher signal-to-noise ratios may be achieved. Forexample, one way to improve the signal-to-noise ratio would be to usemodulation and lock-in techniques. In one embodiment, the light sourcemay be modulated, and then the detection system would be synchronizedwith the light source. In a particular embodiment, the techniques fromlock-in detection may be used, where narrow band filtering around themodulation frequency may be used to reject noise outside the modulationfrequency. In an alternate embodiment, change detection schemes may beused, where the detection system captures the signal with the lightsource on and with the light source off. Again, for this system thelight source may be modulated. Then, the signal with and without thelight source is differenced. This may enable the sun light changes to besubtracted out. In addition, change detection may help to identifyobjects that change in the field of view. In the following someexemplary detection systems are described.

Interference from Skin

Several proposed non-invasive glucose monitoring techniques rely ontransmission, absorption, and/or diffuse reflection through the skin tomeasure blood constituents or blood analytes in veins, arteries,capillaries or in the tissue itself. However, on top of the interferencefrom other blood constituents or analytes, the skin also introducessignificant interference. For example, chemical, structural, andphysiological variations occur that may produce relatively large andnonlinear changes in the optical properties of the tissue sample. In oneembodiment, the near-infrared reflectance or absorbance spectrum may bea complex combination of the tissue scattering properties that resultfrom the concentration and characteristics of a multiplicity of tissuecomponents including water, fat, protein, collagen, elastin, and/orglucose. Moreover, the optical properties of the skin may also changewith environmental factors such as humidity, temperature and pressure.Physiological variation may also cause changes in the tissue measurementover time and may vary based on lifestyle, health, aging, etc. Thestructure and composition of skin may also vary widely amongindividuals, between different sites within an individual, and over timeon the same individual. Thus, the skin introduces a dynamic interferencesignal that may have a wide variation due to a number of parameters.

FIG. 9 shows a schematic cross-section of human skin 900, 901. The toplayer of the skin is epidermis 902, followed by a layer of dermis 903and then subcutaneous fat 904 below the dermis. The epidermis 902, witha thickness of approximately 10-150 microns, may provide a barrier toinfection and loss of moisture and other body constituents. The dermis903 ranges in thickness from approximately 0.5 mm to 4 mm (averagesapproximately 1.2 mm over most of the body) and may provide themechanical strength and elasticity of skin.

In the dermis 903, water may account for approximately 70% of thevolume. The next most abundant constituent in the dermis 903 may becollagen 905, a fibrous protein comprising 70-75% of the dry weight ofthe dermis 903. Elastin fibers 906, also a protein, may also beplentiful in the dermis 903, although they constitute a smaller portionof the bulk. In addition, the dermis 903 may contain a variety ofstructures (e.g., sweat glands, hair follicles with adipose richsebaceous glands 907 near their roots, and blood vessels) and othercellular constituents.

Below the dermis 903 lies the subcutaneous layer 904 comprising mostlyadipose tissue. The subcutaneous layer 904 may be by volumeapproximately 10% water and may be comprised primarily of cells rich intriglycerides or fat. With this complicated structure of the skin 900,901, the concentration of glucose may vary in each layer according to avariety of factors including the water content, the relative sizes ofthe fluid compartments, the distribution of capillaries, the perfusionof blood, the glucose uptake of cells, the concentration of glucose inblood, and the driving forces (e.g., osmotic pressure) behind diffusion.

To better understand the interference that the skin introduces whenattempting to measure glucose, the absorption coefficient for thevarious skin constituents should be examined. For example, FIG. 10illustrates 1000 the absorption coefficients for water (includingscattering) 1001, adipose 1002, collagen 1003 and elastin 1004. Notethat the absorption curves for water 1001 and adipose 1002 arecalibrated, whereas the absorption curves for collagen 1003 and elastin1004 are in arbitrary units. Also shown are vertical lines demarcatingthe wavelengths near 1210 nm 1005 and 1720 nm 1006. In general, thewater absorption increases with increasing wavelength. With theincreasing absorption beyond about 2000 nm, it may be difficult toachieve deeper penetration into biological tissue in the infraredwavelengths beyond approximately 2500 nm.

Although the absorption coefficient may be useful for determining thematerial in which light of a certain infrared wavelength will beabsorbed, to determine the penetration depth of the light of a certainwavelength may also require the addition of scattering loss to thecurves. For example, the water curve 1001 includes the scattering losscurve in addition to the water absorption. In particular, the scatteringloss can be significantly higher at shorter wavelengths. In oneembodiment, near the wavelength of 1720 nm (vertical line 1006 shown inFIG. 10), the adipose absorption 1002 can still be higher than the waterplus scattering loss 1001. For tissue that contains adipose, collagenand elastin, such as the dermis of the skin, the total absorption canexceed the light energy lost to water absorption and light scattering at1720 nm. On the other hand, at 1210 nm the adipose absorption 1002 canbe considerably lower than the water plus scattering loss 1001,particularly since the scattering loss can be dominant at these shorterwavelengths.

The interference for glucose lines observed through skin may beillustrated by overlaying the glucose lines over the absorption curves1000 for the skin constituents. For example, FIG. 2 illustrated that theglucose absorption 202 included features centered around 1587 nm, 1750nm, 2120 nm, 2270 nm and 2320 nm. On FIG. 10 vertical lines have beendrawn at the glucose line wavelengths of 1587 nm 1007, 1750 nm 1008,2120 nm 1009, 2270 nm 1010 and 2320 nm 1011. In one embodiment, it maybe difficult to detect the glucose lines near 1750 nm 1008, 2270 nm 1010and 2320 nm 1011 due to significant spectral interference from otherskin constituents. On the other hand, the glucose line near 1587 m 1007may be more easily detected because it peaks while most of the otherskin constituents are sloped downward toward an absorption valley.Moreover, the glucose line near 2120 nm 1009 may also be detectable forsimilar reasons, although adipose may have conflicting behavior due tolocal absorption minimum and maximum nearby in wavelength.

Thus, beyond the problem of other blood constituents or analytes havingoverlapping spectral features (e.g., FIG. 5), it may be difficult toobserve glucose spectral signatures through the skin and itsconstituents of water, adipose, collagen and elastin. One approach toovercoming this difficulty may be to try to measure the bloodconstituents in veins that are located at relatively shallow distancesbelow the skin. Veins may be more beneficial for the measurement thanarteries, since arteries tend to be located at deeper levels below theskin. Also, in one embodiment it may be advantageous to use adifferential measurement to subtract out some of the interferingabsorption lines from the skin. For example, an instrument head may bedesigned to place one probe above a region of skin over a blood vein,while a second probe may be placed at a region of the skin without anoticeable blood vein below it. Then, by differencing the signals fromthe two probes, at least part of the skin interference may be cancelledout.

Two representative embodiments for performing such a differentialmeasurement are illustrated in FIG. 11 and FIG. 12. In one embodimentshown in FIG. 11, the dorsal of the hand 1100 may be used for measuringblood constituents or analytes. The dorsal of the hand 1100 may haveregions that have distinct veins 1101 as well as regions where the veinsare not as shallow or pronounced 1102. By stretching the hand andleaning it backwards, the veins 1101 may be accentuated in some cases. Anear-infrared diffuse reflectance measurement may be performed byplacing one probe 1103 above the vein-rich region 1101. To turn thisinto a differential measurement, a second probe 1104 may be placed abovea region without distinct veins 1102. Then, the outputs from the twoprobes may be subtracted 1105 to at least partially cancel out thefeatures from the skin. The subtraction may be done preferably in theelectrical domain, although it can also be performed in the opticaldomain or digitally/mathematically using sampled data based on theelectrical and/or optical signals. Although one example of using thedorsal of the hand 1100 is shown, many other parts of the hand can beused within the scope of this disclosure. For example, alternate methodsmay use transmission through the webbing between the thumb and thefingers 1106, or transmission or diffuse reflection through the tips ofthe fingers 1107.

In another embodiment, the dorsal of the foot 1200 may be used insteadof the hand. One advantage of such a configuration may be that forself-testing by a user, the foot may be easier to position theinstrument using both hands. One probe 1203 may be placed over regionswhere there are more distinct veins 1201, and a near-infrared diffusereflectance measurement may be made. For a differential measurement, asecond probe 1204 may be placed over a region with less prominent veins1202, and then the two probe signals may be subtracted, eitherelectronically or optically, or may be digitized/sampled and processedmathematically depending on the particular application andimplementation. As with the hand, the differential measurements may beintended to compensate for or subtract out (at least in part) theinterference from the skin. Since two regions are used in closeproximity on the same body part, this may also aid in removing somevariability in the skin from environmental effects such as temperature,humidity, or pressure. In addition, it may be advantageous to firsttreat the skin before the measurement, by perhaps wiping with a cloth ortreated cotton ball, applying some sort of cream, or placing an ice cubeor chilled bag over the region of interest.

Although two embodiments have been described, many other locations onthe body may be used using a single or differential probe within thescope of this disclosure. In yet another embodiment, the wrist may beadvantageously used, particularly where a pulse rate is typicallymonitored. Since the pulse may be easily felt on the wrist, there isunderlying the region a distinct blood flow. Other embodiments may useother parts of the body, such as the ear lobes, the tongue, the innerlip, the nails, the eye, or the teeth. Some of these embodiments will befurther described below. The ear lobes or the tip of the tongue may beadvantageous because they are thinner skin regions, thus permittingtransmission rather than diffuse reflection. However, the interferencefrom the skin is still a problem in these embodiments. Other regionssuch as the inner lip or the bottom of the tongue may be contemplatedbecause distinct veins are observable, but still the interference fromthe skin may be problematic in these embodiments. The eye may seem as aviable alternative because it is more transparent than skin. However,there are still issues with scattering in the eye. For example, theanterior chamber of the eye (the space between the cornea and the iris)comprises a fluid known as aqueous humor. However, the glucose level inthe eye chamber may have a significant temporal lag on changes in theglucose level compared to the blood glucose level.

Because of the complexity of the interference from skin in non-invasiveglucose monitoring (e.g., FIG. 10), other parts of the body without skinabove blood vessels or capillaries may be alternative candidates formeasuring blood constituents. One embodiment may involve transmission orreflection through human nails. As an example, FIG. 13 illustrates atypical human nail tissue structure 1300 and the capillary vessels belowit. The fingernail 1301 is approximately 1 mm thick, and below thisresides a layer of epidermis 1302 with a thickness of approximately 1mm. The dermis 1304 is also shown, and within particularly the top about0.5 mm of dermis are a significant number of capillary vessels. Tomeasure the blood constituents, the light exposed on the top of thefingernail must penetrate about 2-2.5 mm or more, and the reflectedlight (round trip passage) should be sufficiently strong to measure. Inone embodiment, the distance required to penetrate could be reduced bydrilling a hole in the fingernail 1301.

In this alternative embodiment using the fingernail, there may still beinterference from the nail's spectral features. For example, FIG. 14illustrates the attenuation coefficient 1400 for seven nail samples thatare allowed to stand in an environment with a humidity level of 14%.These coefficients are measured using an FTIR spectrometer over thenear-infrared wavelength range of approximately 1 to 2.5 microns. Thesespectra are believed to correspond to the spectra of keratin containedin the nail plate. The base lines for the different samples are believedto differ because of the influence of scattering. Several of theabsorption peaks observed correspond to peaks of keratin absorption,while other features may appear from the underlying epidermis anddermis. It should also be noted that the attenuation coefficients 1400also vary considerably depending on humidity level or water content aswell as temperature and other environmental factors. Moreover, theattenuation coefficient may also change in the presence of nail polishof various sorts.

Similar to skin, the large variations in attenuation coefficient forfingernails also may interfere with the absorption peaks of glucose. Asan example, in FIG. 14 below the fingernail spectrum is also shown theglucose spectrum 1401 for two different glucose concentrations. Thevertical lines 1402, 1403, 1404, 1405 and 1406 are drawn to illustratethe glucose absorption peaks and where they lie on the fingernailspectra 1400. As is apparent, the nail has interfering features that maybe similar to skin, particularly since both have spectra that vary notonly in wavelength but also with environmental factors. In oneembodiment, it may be possible to see the glucose peaks 1402 and 1404through the fingernail, but it may be much more difficult to observe theglucose peaks near 1403, 1405 and 1406.

Transmission or Reflection Through Teeth

Yet another embodiment may observe the transmittance or reflectancethrough teeth to measure blood constituents or analytes. FIG. 15illustrates an exemplary structure of a tooth 1500. The tooth 1500 has atop layer called the crown 1501 and below that a root 1502 that reacheswell into the gum 1506 and bone 1508 of the mouth. The exterior of thecrown 1501 is an enamel layer 1503, and below the enamel is a layer ofdentine 1504 that sits atop a layer of cementum 1507. Below the dentine1504 is a pulp region 1505, which comprises within it blood vessels 1509and nerves 1510. If the light can penetrate the enamel 1503 and dentine1504, then the blood flow and blood constituents can be measured throughthe blood vessels in the dental pulp 1505. While it may be true that theamount of blood flow in the dental pulp 1505 may be less since itcomprises capillaries, the smaller blood flow could still beadvantageous if there is less interfering spectral features from thetooth.

The transmission, absorption and reflection from teeth has been studiedin the near infrared, and, although there are some features, the enameland dentine appear to be fairly transparent in the near infrared(particularly wavelengths between 1500 and 2500 nm). For example, theabsorption or extinction ratio for light transmission has been studied.FIG. 16A illustrates the attenuation coefficient 1600 for dental enamel1601 (filled circles) and the absorption coefficient of water 1602 (opencircles) versus wavelength. Near-infrared light may penetrate muchfurther without scattering through all the tooth enamel, due to thereduced scattering coefficient in normal enamel. Scattering in enamelmay be fairly strong in the visible, but decreases as approximately1/wavelength³ (i.e., inverse of the wavelength cubed) with increasingwavelength to a value of only 2-3 cm⁻¹ at 1310 nm and 1550 nm in thenear infrared. Therefore, enamel may be virtually transparent in thenear infrared with optical attenuation 1-2 orders of magnitude less thanin the visible range.

As another example, FIG. 16B illustrates the absorption spectrum 1650 ofintact enamel 1651 (dashed line) and dentine 1652 (solid line) in thewavelength range of approximately 1.2 to 2.4 microns. In the nearinfrared there are two absorption bands around 1.5 and 2 microns. Theband with a peak around 1.57 microns may be attributed to the overtoneof valent vibration of water present in both enamel and dentine. In thisband, the absorption is greater for dentine than for enamel, which maybe related to the large water content in this tissue. In the region of 2microns, dentine may have two absorption bands, and enamel one. The bandwith a maximum near 2.1 microns may belong to the overtone of vibrationof PO hydroxyapatite groups, which is the main substance of both enameland dentine. Moreover, the band with a peak near 1.96 microns in dentinemay correspond to water absorption (dentine may contain substantiallyhigher water than enamel).

In addition to the absorption coefficient, the reflectance from intactteeth and teeth with dental caries (e.g., cavities) has been studied. Inone embodiment, FIG. 17 shows the near infrared spectral reflectance1700 over the wavelength range of approximately 800 nm to 2500 nm froman occlusal (e.g., top/bottom) tooth surface 1704. The curve with blackdiamonds 1701 corresponds to the reflectance from a sound, intact toothsection. The curve with asterisks * 1702 corresponds to a tooth sectionwith an enamel lesion. The curve with circles 1703 corresponds to atooth section with a dentine lesion. Thus, when there is a lesion, morescattering occurs and there may be an increase in the reflected light.

For wavelengths shorter than approximately 1400 nm, the shapes of thespectra remain similar, but the amplitude of the reflection changes withlesions. Between approximately 1400 nm and 2500 nm, an intact tooth 1701has low reflectance (e.g., high transmission), and the reflectanceappears to be more or less independent of wavelength. On the other hand,in the presence of lesions 1702 and 1703, there is increased scattering,and the scattering loss may be wavelength dependent. For example, thescattering loss may decrease as 1/(wavelength)³—so, the scattering lossdecreases with longer wavelengths. When there is a lesion in the dentine1703, more water can accumulate in the area, so there is also increasedwater absorption. For example, the dips near 1450 nm and 1900 nmcorrespond to water absorption, and the reflectance dips areparticularly pronounced in the dentine lesion 1703. One other benefit ofthe absorption, transmission or reflectance in the near infrared may bethat stains and non-calcified plaque are not visible in this wavelengthrange, enabling better discrimination of defects, cracks, anddemineralized areas.

Compared with the interference from skin 1000 in FIG. 10 or fingernails1400 in FIG. 14, the teeth appear to introduce much less interferencefor non-invasive monitoring of blood constituents. The few features inFIG. 16B or 17 may be calibrated out of the measurement. Also, using anintact tooth 1701 may further minimize any interfering signals.Furthermore, since the tooth comprises relatively hard tissue, higherpower from the light sources in the near infrared may be used withoutdamaging the tissue, such as with skin.

Human Interface for Measurement System

A number of different types of measurements may be used to sample theblood in the dental pulp. The basic feature of the measurements shouldbe that the optical properties are measured as a function of wavelengthat a plurality of wavelengths. As further described below, the lightsource may output a plurality of wavelengths, or a continuous spectrumover a range of wavelengths. In a preferred embodiment, the light sourcemay cover some or all of the wavelength range between approximately 1400nm and 2500 nm. The signal may be received at a receiver, which may alsocomprise a spectrometer or filters to discriminate between differentwavelengths. The signal may also be received at a camera, which may alsocomprise filters or a spectrometer. In an alternate embodiment, thespectral discrimination using filters or a spectrometer may be placedafter the light source rather than at the receiver. The receiver usuallycomprises one or more detectors (optical-to-electrical conversionelement) and electrical circuitry. The receiver may also be coupled toanalog to digital converters, particularly if the signal is to be fed toa digital device.

Referring to FIG. 15, one or more light sources 1511 may be used forillumination. In one embodiment, a transmission measurement may beperformed by directing the light source output 1511 to the region nearthe interface between the gum 1506 and dentine 1504. In one embodiment,the light may be directed using a light guide or a fiber optic. Thelight may then propagate through the dental pulp 1505 to the other side,where the light may be incident on one or more detectors or anotherlight guide to transport the signal to a spectrometer, receiver orcamera 1512. In another embodiment, the light source may be directed toone or more locations near the interface between the gum 1506 anddentine 1504 (in one example, could be from the two sides of the tooth).The transmitted light may then be detected in the occlusal surface abovethe tooth using a spectrometer, receiver, or camera 1512. In yet anotherembodiment, a reflectance measurement may be conducted by directing thelight source output 1511 to, for example, the occlusal surface of thetooth, and then detecting the reflectance at a spectrometer, receiver orcamera 1513. Although a few embodiments for measuring the bloodconstituents through a tooth are described, other embodiments andtechniques may also be used and are intended to be covered by thisdisclosure.

The human interface for the non-invasive measurement of bloodconstituents may be of various forms. In one embodiment, a “clamp”design 1800 may be used cap over one or more teeth, as illustrated inFIG. 18A. The clamp design may be different for different types ofteeth, or it may be flexible enough to fit over different types ofteeth. For example, different types of teeth include the molars (towardthe back of the mouth), the premolars, the canine, and the incisors(toward the front of the mouth). One embodiment of the clamp-type designis illustrated in FIG. 18A for a molar tooth 1808. The C-clamp 1801 maybe made of a plastic or rubber material, and it may comprise a lightsource input 1802 and a detector output 1803 on the front or back of thetooth.

The light source input 1802 may comprise a light source directly, or itmay have light guided to it from an external light source. Also, thelight source input 1802 may comprise a lens system to collimate or focusthe light across the tooth. The detector output 1803 may comprise adetector directly, or it may have a light guide to transport the signalto an external detector element. The light source input 1802 may becoupled electrically or optically through 1804 to a light input 1806.For example, if the light source is external in 1806, then the couplingelement 1804 may be a light guide, such as a fiber optic. Alternately,if the light source is contained in 1802, then the coupling element 1804may be electrical wires connecting to a power supply in 1806. Similarly,the detector output 1803 may be coupled to a detector output unit 1807with a coupling element 1805, which may be one or more electrical wiresor a light guide, such as a fiber optic. This is just one example of aclamp over one or more teeth, but other embodiments may also be used andare intended to be covered by this disclosure.

In yet another embodiment, one or more light source ports and sensorports may be used in a mouth-guard type design. For example, oneembodiment of a dental mouth guard 1850 is illustrated in FIG. 18B. Thestructure of the mouth guard 1851 may be similar to mouth guards used insports (e.g., when playing football or boxing) or in dental trays usedfor applying fluoride treatment, and the mouth guard may be made fromplastic or rubber materials, for example. As an example, the mouth guardmay have one or more light source input ports 1852, 1853 and one or moredetector output ports 1854, 1855. Although six input and output portsare illustrated, any number of ports may be used.

Similar to the clamp design described above, the light source inputs1852, 1853 may comprise one or more light sources directly, or they mayhave light guided to them from an external light source. Also, the lightsource inputs 1852, 1853 may comprise lens systems to collimate or focusthe light across the teeth. The detector outputs 1854, 1855 may compriseone or more detectors directly, or they may have one or more lightguides to transport the signals to an external detector element. Thelight source inputs 1852, 1853 may be coupled electrically or opticallythrough 1856 to a light input 1857. For example, if the light source isexternal in 1857, then the one or more coupling elements 1856 may be oneor more light guides, such as a fiber optic. Alternately, if the lightsources are contained in 1852, 1853, then the coupling element 1856 maybe one or more electrical wires connecting to a power supply in 1857.Similarly, the detector outputs 1854, 1855 may be coupled to a detectoroutput unit 1859 with one or more coupling elements 1858, which may beone or more electrical wires or one or more light guides, such as afiber optic. This is just one example of a mouth guard design covering aplurality of teeth, but other embodiments may also be used and areintended to be covered by this disclosure. For instance, the position ofthe light source inputs and detector output ports could be exchanged, orsome mixture of locations of light source inputs and detector outputports could be used.

Also, if reflectance from the teeth is to be measured, then the lightsources and detectors may be on the same side of the tooth. Moreover, itmay be advantageous to pulse the light source with a particular pulsewidth and pulse repetition rate, and then the detection system canmeasure the pulsed light returned from or transmitted through the tooth.Using a lock-in type technique (e.g., detecting at the same frequency asthe pulsed light source and also possibly phase locked to the samesignal), the detection system may be able to reject background orspurious signals and increase the signal-to-noise ratio of themeasurement.

Other elements may be added to the human interface designs of FIG. 18and are also intended to be covered by this disclosure. For instance, inone embodiment it may be desirable to have replaceable inserts that maybe disposable. Particularly in a doctor's office or hospital setting,the same instrument may be used with a plurality of patients. Ratherthan disinfecting the human interface after each use, it may bepreferable to have disposable inserts that can be thrown away after eachuse. In one embodiment, a thin plastic coating material may enclose theclamp design of FIG. 18A or mouth guard design of FIG. 18B. The coatingmaterial may be inserted before each use, and then after the measurementis exercised the coating material may be peeled off and replaced. Such adesign may save the physician or user considerable time, while at thesame time provide the business venture with a recurring cost revenuesource. Any coating material or other disposable device may beconstructed of a material having suitable optical properties that may beconsidered during processing of the signals used to detect any anomaliesin the teeth.

Light Sources for Near Infrared

In general, the near-infrared (NIR) region of the electromagneticspectrum covers between approximately 0.7 microns (700 nm) to about 2.5microns (2500 nm). However, it may also be advantageous to use just theshort-wave infrared between approximately 1.4 microns (1400 nm) andabout 2.5 microns (2500 nm). One reason for preferring the SWIR over theentire NIR may be to operate in the so-called “eye-safe” window, whichcorresponds to wavelengths longer than about 1400 nm. While the SWIR isused for illustrative purposes, it should be clear that the discussionthat follows could also apply to using the NIR wavelength range, orother wavelength bands. There are a number of light sources that may beused in the near infrared. To be more specific, the discussion belowwill consider light sources operating in the so-called short waveinfrared (SWIR), which may cover the wavelength range of approximately1400 nm to 2500 nm. Other wavelength ranges may also be used for theapplications described in this disclosure, so the discussion below ismerely provided for exemplary types of light sources. The SWIRwavelength range may be valuable for a number of reasons. First, theSWIR corresponds to a transmission window through water and theatmosphere. For example, 302 in FIG. 3A and 1602 in FIG. 16A illustratethe water transmission windows. Also, through the atmosphere,wavelengths in the SWIR have similar transmission windows due to watervapor in the atmosphere. Second, the so-called “eye-safe” wavelengthsare wavelengths longer than approximately 1400 nm. Third, the SWIRcovers the wavelength range for nonlinear combinations of stretching andbending modes as well as the first overtone of C—H stretching modes.Thus, for example, glucose and ketones among other substances may haveunique signatures in the SWIR. Moreover, many solids have distinctspectral signatures in the SWIR, so particular solids may be identifiedusing stand-off detection or remote sensing. For instance, manyexplosives have unique signatures in the SWIR.

Infrared light sources can be used for diagnostics and therapeutics in anumber of medical applications. For example, broadband light sources canadvantageously be used for diagnostics, while narrower band lightsources can advantageously be used for therapeutics. In one embodiment,selective absorption or damage can be achieved by choosing the laserwavelength to lie approximately at an absorption peak of particulartissue types. Also, by using infrared wavelengths that minimize waterabsorption peaks and longer wavelengths that have lower tissuescattering, larger penetration depths into the biological tissue can beobtained. In this disclosure, infrared wavelengths include wavelengthsin the range of approximately 0.9 microns to 10 microns, withwavelengths between about 0.98 microns and 2.5 microns more suitable forcertain applications.

Different light sources may be selected for the SWIR based on the needsof the application. Some of the features for selecting a particularlight source include power or intensity, wavelength range or bandwidth,spatial or temporal coherence, spatial beam quality for focusing ortransmission over long distance, and pulse width or pulse repetitionrate. Depending on the application, lamps, light emitting diodes (LEDs),laser diodes (LD's), tunable LD's, super-luminescent laser diodes(SLDs), fiber lasers or super-continuum sources (SC) may beadvantageously used. Also, different fibers may be used for transportingthe light, such as fused silica fibers, plastic fibers, mid-infraredfibers (e.g., tellurite, chalcogenides, fluorides, ZBLAN, etc), or ahybrid of these fibers.

Lamps may be used if low power or intensity of light is required in theSWIR, and if an incoherent beam is suitable. In one embodiment, in theSWIR an incandescent lamp that can be used is based on tungsten andhalogen, which have an emission wavelength between approximately 500 nmto 2500 nm. For low intensity applications, it may also be possible touse thermal sources, where the SWIR radiation is based on the black bodyradiation from the hot object. Although the thermal and lamp basedsources are broadband and have low intensity fluctuations, it may bedifficult to achieve a high signal-to-noise ratio in a non-invasiveblood constituent measurement due to the low power levels. Also, thelamp based sources tend to be energy inefficient.

In another embodiment, LED's can be used that have a higher power levelin the SWIR wavelength range. LED's also produce an incoherent beam, butthe power level can be higher than a lamp and with higher energyefficiency. Also, the LED output may more easily be modulated, and theLED provides the option of continuous wave or pulsed mode of operation.LED's are solid state components that emit a wavelength band that is ofmoderate width, typically between about 20 nm to 40 nm. There are alsoso-called super-luminescent LEDs that may even emit over a much widerwavelength range. In another embodiment, a wide band light source may beconstructed by combining different LEDs that emit in differentwavelength bands, some of which could preferably overlap in spectrum.One advantage of LEDs as well as other solid state components is thecompact size that they may be packaged into.

In yet another embodiment, various types of laser diodes may be used inthe SWIR wavelength range. Just as LEDs may be higher in power butnarrower in wavelength emission than lamps and thermal sources, the LDsmay be yet higher in power but yet narrower in wavelength emission thanLEDs. Different kinds of LDs may be used, including Fabry-Perot LDs,distributed feedback (DFB) LDs, distributed Bragg reflector (DBR) LDs.Since the LDs have relatively narrow wavelength range (typically under10 nm), in one embodiment a plurality of LDs may be used that are atdifferent wavelengths in the SWIR. For example, in a preferredembodiment for non-invasive glucose monitoring, it may be advantageousto use LDs having emission spectra near some or all of the glucosespectral peaks (e.g., near 1587 nm, 1750 nm, 2120 nm, 2270 nm, and 2320nm). The various LDs may be spatially multiplexed, polarizationmultiplexed, wavelength multiplexed, or a combination of thesemultiplexing methods. Also, the LDs may be fiber pig-tailed or have oneor more lenses on the output to collimate or focus the light. Anotheradvantage of LDs is that they may be packaged compactly and may have aspatially coherent beam output. Moreover, tunable LDs that can tune overa range of wavelengths are also available. The tuning may be done byvarying the temperature, or electrical current may be used in particularstructures, such as distributed Bragg reflector LDs. In anotherembodiment, external cavity LDs may be used that have a tuning element,such as a fiber grating or a bulk grating, in the external cavity.

In another embodiment, super-luminescent laser diodes may provide higherpower as well as broad bandwidth. An SLD is typically an edge emittingsemiconductor light source based on super-luminescence (e.g., this couldbe amplified spontaneous emission). SLDs combine the higher power andbrightness of LDs with the low coherence of conventional LEDs, and theemission band for SLD's may be 5 to 100 nm wide, preferably in the 60 to100 nm range. Although currently SLDs are commercially available in thewavelength range of approximately 400 nm to 1700 nm, SLDs could and mayin the future be made to cover a broader region of the SWIR.

In yet another embodiment, high power LDs for either direct excitationor to pump fiber lasers and SC light sources may be constructed usingone or more laser diode bar stacks. As an example, FIG. 19 shows anexample of the block diagram 1900 or building blocks for constructingthe high power LDs. In this embodiment, one or more diode bar stacks1901 may be used, where the diode bar stack may be an array of severalsingle emitter LDs. Since the fast axis (e.g., vertical direction) maybe nearly diffraction limited while the slow-axis (e.g., horizontalaxis) may be far from diffraction limited, different collimators 1902may be used for the two axes.

Then, the brightness may be increased by spatially combining the beamsfrom multiple stacks 1903. The combiner may include spatialinterleaving, it may include wavelength multiplexing, or it may involvea combination of the two. Different spatial interleaving schemes may beused, such as using an array of prisms or mirrors with spacers to bendone array of beams into the beam path of the other. In anotherembodiment, segmented mirrors with alternate high-reflection andanti-reflection coatings may be used. Moreover, the brightness may beincreased by polarization beam combining 1904 the two orthogonalpolarizations, such as by using a polarization beam splitter. In oneembodiment, the output may then be focused or coupled into a largediameter core fiber. As an example, typical dimensions for the largediameter core fiber range from approximately 100 microns in diameter to400 microns or more. Alternatively or in addition, a custom beam shapingmodule 1905 may be used, depending on the particular application. Forexample, the output of the high power LD may be used directly 1906, orit may be fiber coupled 1907 to combine, integrate, or transport thehigh power LD energy. These high power LDs may grow in importancebecause the LD powers can rapidly scale up. For example, instead of thepower being limited by the power available from a single emitter, thepower may increase in multiples depending on the number of diodesmultiplexed and the size of the large diameter fiber. Although FIG. 19is shown as one embodiment, some or all of the elements may be used in ahigh power LD, or additional elements may also be used.

SWIR Super-Continuum Lasers

Each of the light sources described above have particular strengths, butthey also may have limitations. For example, there is typically atrade-off between wavelength range and power output. Also, sources suchas lamps, thermal sources, and LEDs produce incoherent beams that may bedifficult to focus to a small area and may have difficulty propagatingfor long distances. An alternative source that may overcome some ofthese limitations is an SC light source. Some of the advantages of theSC source may include high power and intensity, wide bandwidth,spatially coherent beam that can propagate nearly transform limited overlong distances, and easy compatibility with fiber delivery.

Supercontinuum lasers may combine the broadband attributes of lamps withthe spatial coherence and high brightness of lasers. By exploiting amodulational instability initiated supercontinuum (SC) mechanism, anall-fiber-integrated SC laser with no moving parts may be built usingcommercial-off-the-shelf (COTS) components. Moreover, the fiber laserarchitecture may be a platform where SC in the visible,near-infrared/SWIR, or mid-IR can be generated by appropriate selectionof the amplifier technology and the SC generation fiber. But until now,SC lasers were used primarily in laboratory settings since typicallylarge, table-top, mode-locked lasers were used to pump nonlinear mediasuch as optical fibers to generate SC light. However, those large pumplasers may now be replaced with diode lasers and fiber amplifiers thatgained maturity in the telecommunications industry.

In one embodiment, an all-fiber-integrated, high-powered SC light source2000 may be elegant for its simplicity (FIG. 20). The light may be firstgenerated from a seed laser diode 2001. For example, the seed LD 2001may be a distributed feedback laser diode with a wavelength near 1542 or1550 nm, with approximately 0.5-2.0 ns pulsed output, and with a pulserepetition rate between a kilohertz to about 100 MHz or more. The outputfrom the seed laser diode may then be amplified in a multiple-stagefiber amplifier 2002 comprising one or more gain fiber segments. In oneembodiment, the first stage pre-amplifier 2003 may be designed foroptimal noise performance. For example, the pre-amplifier 2003 may be astandard erbium-doped fiber amplifier or an erbium/ytterbium dopedcladding pumped fiber amplifier. Between amplifier stages 2003 and 2006,it may be advantageous to use band-pass filters 2004 to block amplifiedspontaneous emission and isolators 2005 to prevent spurious reflections.Then, the power amplifier stage 2006 may use a cladding-pumped fiberamplifier that may be optimized to minimize nonlinear distortion. Thepower amplifier fiber 2006 may also be an erbium-doped fiber amplifier,if only low or moderate power levels are to be generated.

The SC generation 2007 may occur in the relatively short lengths offiber that follow the pump laser. In one exemplary embodiment, the SCfiber length may range from a few millimeters to 100 m or more. In oneembodiment, the SC generation may occur in a first fiber 2008 where themodulational-instability initiated pulse break-up primarily occurs,followed by a second fiber 2009 where the SC generation and spectralbroadening primarily occurs.

In one embodiment, one or two meters of standard single-mode fiber (SMF)after the power amplifier stage may be followed by several meters of SCgeneration fiber. For this example, in the SMF the peak power may beseveral kilowatts and the pump light may fall in the anomalousgroup-velocity dispersion regime—often called the soliton regime. Forhigh peak powers in the dispersion regime, the nanosecond pulses may beunstable due to a phenomenon known as modulational instability, which isbasically parametric amplification in which the fiber nonlinearity helpsto phase match the pulses. As a consequence, the nanosecond pump pulsesmay be broken into many shorter pulses as the modulational instabilitytries to form soliton pulses from the quasi-continuous-wave background.Although the laser diode and amplification process starts withapproximately nanosecond-long pulses, modulational instability in theshort length of SMF fiber may form approximately 0.5 ps toseveral-picosecond-long pulses with high intensity. Thus, the few metersof SMF fiber may result in an output similar to that produced bymode-locked lasers, except in a much simpler and cost-effective manner.

The short pulses created through modulational instability may then becoupled into a nonlinear fiber for SC generation. The nonlinearmechanisms leading to broadband SC may include four-wave mixing orself-phase modulation along with the optical Raman effect. Since theRaman effect is self-phase-matched and shifts light to longerwavelengths by emission of optical photons, the SC may spread to longerwavelengths very efficiently. The short-wavelength edge may arise fromfour-wave mixing, and often times the short wavelength edge may belimited by increasing group-velocity dispersion in the fiber. In manyinstances, if the particular fiber used has sufficient peak power and SCfiber length, the SC generation process may fill the long-wavelengthedge up to the transmission window.

Mature fiber amplifiers for the power amplifier stage 2006 includeytterbium-doped fibers (near 1060 nm), erbium-doped fibers (near 1550nm), erbium/ytterbium-doped fibers (near 1550 nm), or thulium-dopedfibers (near 2000 nm). In various embodiments, candidates for SC fiber2009 include fused silica fibers (for generating SC between 0.8-2.7 μm),mid-IR fibers such as fluorides, chalcogenides, or tellurites (forgenerating SC out to 4.5 μm or longer), photonic crystal fibers (forgenerating SC between 0.4 and 1.7 μm), or combinations of these fibers.Therefore, by selecting the appropriate fiber-amplifier doping for 2006and nonlinear fiber 2009, SC may be generated in the visible,near-IR/SWIR, or mid-IR wavelength region.

The configuration 2000 of FIG. 20 is just one particular example, andother configurations can be used and are intended to be covered by thisdisclosure. For example, further gain stages may be used, and differenttypes of lossy elements or fiber taps may be used between the amplifierstages. In another embodiment, the SC generation may occur partially inthe amplifier fiber and in the pig-tails from the pump combiner or otherelements. In yet another embodiment, polarization maintaining fibers maybe used, and a polarizer may also be used to enhance the polarizationcontrast between amplifier stages. Also, not discussed in detail aremany accessories that may accompany this set-up, such as driverelectronics, pump laser diodes, safety shut-offs, and thermal managementand packaging.

One example of an SC laser that operates in the SWIR used in oneembodiment is illustrated in FIG. 21. This SWIR SC source 2100 producesan output of up to approximately 5 W over a spectral range of about 1.5to 2.4 microns, and this particular laser is made out of polarizationmaintaining components. The seed laser 2101 is a distributed feedback(DFB) laser operating near 1542 nm producing approximately 0.5nanosecond (ns) pulses at an about 8 MHz repetition rate. Thepre-amplifier 2102 is forward pumped and uses about 2 m length oferbium/ytterbium cladding pumped fiber 2103 (often also called dual-corefiber) with an inner core diameter of 12 microns and outer core diameterof 130 microns. The pre-amplifier gain fiber 2103 is pumped using a 10 W940 nm laser diode 2105 that is coupled in using a fiber combiner 2104.

In this particular 5 W unit, the mid-stage between amplifier stages 2102and 2106 comprises an isolator 2107, a band-pass filter 2108, apolarizer 2109 and a fiber tap 2110. The power amplifier 2106 uses a 4 mlength of the 12/130 micron erbium/ytterbium doped fiber 2111 that iscounter-propagating pumped using one or more 30 W 940 nm laser diodes2112 coupled in through a combiner 2113. An approximately 1-2 meterlength of the combiner pig-tail helps to initiate the SC process, andthen a length of PM-1550 fiber 2115 (polarization maintaining,single-mode, fused silica fiber optimized for 1550 nm) is spliced 2114to the combiner output.

If an output fiber of about 10 m in length is used, then the resultingoutput spectrum 2200 is shown in FIG. 22. The details of the outputspectrum 2200 depend on the peak power into the fiber, the fiber length,and properties of the fiber such as length and core size, as well as thezero dispersion wavelength and the dispersion properties. For example,if a shorter length of fiber is used, then the spectrum actually reachesto longer wavelengths (e.g., a 2 m length of SC fiber broadens thespectrum to −2500 nm). Also, if extra-dry fibers are used with less O—Hcontent, then the wavelength edge may also reach to a longer wavelength.To generate more spectrum toward the shorter wavelengths, the pumpwavelength (in this case .about.1542 nm) should be close to the zerodispersion wavelength in the fiber. For example, by using a dispersionshifted fiber or so-called non-zero dispersion shifted fiber, the shortwavelength edge may shift to shorter wavelengths.

Although one particular example of a 5 W SWIR-SC has been described,different components, different fibers, and different configurations mayalso be used consistent with this disclosure. For instance, anotherembodiment of the similar configuration 2100 in FIG. 21 may be used togenerate high powered SC between approximately 1060 and 1800 nm. Forthis embodiment, the seed laser 2101 may be a 1064 nm distributedfeedback (DFB) laser diode, the pre-amplifier gain fiber 2103 may be aytterbium-doped fiber amplifier with 10/125 microns dimensions, and thepump laser 2105 may be a 10 W 915 nm laser diode. In the mid-stage, amode field adapter may be included in addition to the isolator 2107,band pass filter 2108, polarizer 2109 and tap 2110. The gain fiber 2111in the power amplifier may be a 20 m length of ytterbium-doped fiberwith 25/400 microns dimension for example. The pump 2112 for the poweramplifier may be up to six pump diodes providing 30 W each near 915 nm,for example. For this much pump power, the output power in the SC may beas high as 50 W or more.

In another embodiment, it may be desirous to generate high power SWIR SCover 1.4-1.8 microns and separately 2-2.5 microns (the window between1.8 and 2 microns may be less important due to the strong water andatmospheric absorption). For example, the top SC source of FIG. 23 canlead to bandwidths ranging from about 1400 nm to 1800 nm or broader,while the lower SC source of FIG. 23 can lead to bandwidths ranging fromabout 1900 nm to 2500 nm or broader. Since these wavelength ranges areshorter than about 2500 nm, the SC fiber can be based on fused silicafiber. Exemplary SC fibers include standard single-mode fiber SMF,high-nonlinearity fiber, high-NA fiber, dispersion shifted fiber,dispersion compensating fiber, and photonic crystal fibers.Non-fused-silica fibers can also be used for SC generation, includingchalcogenides, fluorides, ZBLAN, tellurites, and germanium oxide fibers.

In one embodiment, the top of FIG. 23 illustrates a block diagram for anSC source 2300 capable of generating light between approximately 1400and 1800 nm or broader. As an example, a pump fiber laser similar toFIG. 21 can be used as the input to a SC fiber 2309. The seed laserdiode 2301 can comprise a DFB laser that generates, for example, severalmilliwatts of power around 1542 or 1553 nm. The fiber pre-amplifier 2302can comprise an erbium-doped fiber amplifier or an erbium/ytterbiumdoped double clad fiber. In this example a mid-stage amplifier 2303 canbe used, which can comprise an erbium/ytterbium doped double-clad fiber.A bandpass filter 2305 and isolator 2306 may be used between thepre-amplifier 2302 and mid-stage amplifier 2303. The power amplifierstage 2304 can comprise a larger core size erbium/ytterbium dopeddouble-clad fiber, and another bandpass filter 2307 and isolator 2308can be used before the power amplifier 2304. The output of the poweramplifier can be coupled to the SC fiber 2309 to generate the SC output2310. This is just one exemplary configuration for an SC source, andother configurations or elements may be used consistent with thisdisclosure.

In yet another embodiment, the bottom of FIG. 23 illustrates a blockdiagram for an SC source 2350 capable of generating light betweenapproximately 1900 and 2500 nm or broader. As an example, the seed laserdiode 2351 can comprise a DFB or DBR laser that generates, for example,several milliwatts of power around 1542 or 1553 nm. The fiberpre-amplifier 2352 can comprise an erbium-doped fiber amplifier or anerbium/ytterbium doped double-clad fiber. In this example a mid-stageamplifier 2353 can be used, which can comprise an erbium/ytterbium dopeddouble-clad fiber. A bandpass filter 2355 and isolator 2356 may be usedbetween the pre-amplifier 2352 and mid-stage amplifier 2353. The poweramplifier stage 2354 can comprise a thulium doped double-clad fiber, andanother isolator 2357 can be used before the power amplifier 2354. Notethat the output of the mid-stage amplifier 2353 can be approximatelynear 1550 nm, while the thulium-doped fiber amplifier 2354 can amplifywavelengths longer than approximately 1900 nm and out to about 2100 nm.Therefore, for this configuration wavelength shifting may be requiredbetween 2353 and 2354. In one embodiment, the wavelength shifting can beaccomplished using a length of standard single-mode fiber 2358, whichcan have a length between approximately 5 and 50 meters, for example.The output of the power amplifier 2354 can be coupled to the SC fiber2359 to generate the SC output 2360. This is just one exemplaryconfiguration for an SC source, and other configurations or elements canbe used consistent with this disclosure. For example, the variousamplifier stages can comprise different amplifier types, such as erbiumdoped fibers, ytterbium doped fibers, erbium/ytterbium co-doped fibersand thulium doped fibers. One advantage of the SC lasers illustrated inFIGS. 20-23 are that they may use all-fiber components, so that the SClaser can be all-fiber, monolithically integrated with no moving parts.The all-integrated configuration can consequently be robust andreliable.

FIGS. 20-23 are examples of SC light sources that may be advantageouslyused for SWIR light generation in various medical diagnostic andtherapeutic applications. However, many other versions of the SC lightsources may also be made that are intended to also be covered by thisdisclosure. For example, the SC generation fiber could be pumped by amode-locked laser, a gain-switched semiconductor laser, an opticallypumped semiconductor laser, a solid state laser, other fiber lasers, ora combination of these types of lasers. Also, rather than using a fiberfor SC generation, either a liquid or a gas cell might be used as thenonlinear medium in which the spectrum is to be broadened.

Even within the all-fiber versions illustrated such as in FIG. 21,different configurations could be used consistent with the disclosure.In an alternate embodiment, it may be desirous to have a lower costversion of the SWIR SC laser of FIG. 21. One way to lower the cost couldbe to use a single stage of optical amplification, rather than twostages, which may be feasible if lower output power is required or thegain fiber is optimized. For example, the pre-amplifier stage 2102 mightbe removed, along with at least some of the mid-stage elements. In yetanother embodiment, the gain fiber could be double passed to emulate atwo stage amplifier. In this example, the pre-amplifier stage 2102 mightbe removed, and perhaps also some of the mid-stage elements. A mirror orfiber grating reflector could be placed after the power amplifier stage2106 that may preferentially reflect light near the wavelength of theseed laser 2101. If the mirror or fiber grating reflector can transmitthe pump light near 940 nm, then this could also be used instead of thepump combiner 2113 to bring in the pump light 2112. The SC fiber 2115could be placed between the seed laser 2101 and the power amplifierstage 2106 (SC is only generated after the second pass through theamplifier, since the power level may be sufficiently high at that time).In addition, an output coupler may be placed between the seed laserdiode 2101 and the SC fiber, which now may be in front of the poweramplifier 2106. In a particular embodiment, the output coupler could bea power coupler or divider, a dichroic coupler (e.g., passing seed laserwavelength but outputting the SC wavelengths), or a wavelength divisionmultiplexer coupler. This is just one further example, but a myriad ofother combinations of components and architectures could also be usedfor SC light sources to generate SWIR light that are intended to becovered by this disclosure.

Wireless Link to the Cloud

The non-invasive blood constituent or analytes measurement device mayalso benefit from communicating the data output to the “cloud” (e.g.,data servers and processors in the web remotely connected) via wiredand/or wireless communication strategies. The non-invasive devices maybe part of a series of biosensors applied to the patient, andcollectively these devices form what might be called a body area networkor a personal area network. The biosensors and non-invasive devices maycommunicate to a smart phone, tablet, personal data assistant, computer,and/or other microprocessor-based device, which may in turn wirelesslyor over wire and/or fiber optically transmit some or all of the signalor processed data to the internet or cloud. The cloud or internet may inturn send the data to doctors or health care providers as well as thepatients themselves. Thus, it may be possible to have a panoramic,high-definition, relatively comprehensive view of a patient that doctorscan use to assess and manage disease, and that patients can use to helpmaintain their health and direct their own care.

In a particular embodiment 2400, the physiological measurement device ornon-invasive blood constituent measurement device 2401 may comprise atransmitter 2403 to communicate over a first communication link 2404 inthe body area network or personal area network to a receiver in a smartphone, tablet cell phone, PDA, or computer 2405. For the measurementdevice 2401, it may also be advantageous to have a processor 2402 toprocess some of the physiological data, since with processing the amountof data to transmit may be less (hence, more energy efficient). Thefirst communication link 2404 may operate through the use of one of manywireless technologies such as Bluetooth, Zigbee, WiFi, IrDA (infrareddata association), wireless USB, or Z-wave, to name a few.Alternatively, the communication link 2404 may occur in the wirelessmedical band between 2360 and 2390 MHz, which the FCC allocated formedical body area network devices, or in other designated medical deviceor WMTS bands. These are examples of devices that can be used in thebody area network and surroundings, but other devices could also be usedand are included in the scope of this disclosure.

The personal device 2405 may store, process, display, and transmit someof the data from the measurement device 2401. The device 2405 maycomprise a receiver, transmitter, display, voice control and speakers,and one or more control buttons or knobs and a touch screen. Examples ofthe device 2405 include smart phones such as the Apple iPhones® orphones operating on the Android or Microsoft systems. In one embodiment,the device 2405 may have an application, software program, or firmwareto receive and process the data from the measurement device 2401. Thedevice 2405 may then transmit some or all of the data or the processeddata over a second communication link 2406 to the internet or “cloud”2407. The second communication link 2406 may advantageously comprise atleast one segment of a wireless transmission link, which may operateusing WiFi or the cellular network. The second communication link 2406may additionally comprise lengths of fiber optic and/or communicationover copper wires or cables.

The internet or cloud 2407 may add value to the measurement device 2401by providing services that augment the physiological data collected. Ina particular embodiment, some of the functions performed by the cloudinclude: (a) receive at least a fraction of the data from the device2405; (b) buffer or store the data received; (c) process the data usingsoftware stored on the cloud; (d) store the resulting processed data;and (e) transmit some or all of the data either upon request or based onan alarm. As an example, the data or processed data may be transmitted2408 back to the originator (e.g., patient or user), it may betransmitted 2409 to a health care provider or doctor, or it may betransmitted 2410 to other designated recipients.

The cloud 2407 may provide a number of value-add services. For example,the cloud application may store and process the physiological data forfuture reference or during a visit with the healthcare provider. If apatient has some sort of medical mishap or emergency, the physician canobtain the history of the physiological parameters over a specifiedperiod of time. In another embodiment, if the physiological parametersfall out of acceptable range, alarms may be delivered to the user 2408,the healthcare provider 2409, or other designated recipients 2410. Theseare just some of the features that may be offered, but many others maybe possible and are intended to be covered by this disclosure. As anexample, the device 2405 may also have a GPS sensor, so the cloud 2407may be able to provide time, data and position along with thephysiological parameters. Thus, if there is a medical emergency, thecloud 2407 could provide the location of the patient to the healthcareprovider 2409 or other designated recipients 2410. Moreover, thedigitized data in the cloud 2407 may help to move toward what is oftencalled “personalized medicine.” Based on the physiological parameterdata history, medication or medical therapies may be prescribed that arecustomized to the particular patient.

Beyond the above benefits, the cloud application 2407 and application onthe device 2405 may also have financial value for companies developingmeasurement devices 2401 such as a non-invasive blood constituentmonitor. In the case of glucose monitors, the companies make themajority of their revenue on the measurement strips. However, with anon-invasive monitor, there is no need for strips, so there is less ofan opportunity for recurring costs (e.g., the razor/razor blade modeldoes not work for non-invasive devices). On the other hand, people maybe willing to pay a periodic fee for the value-add services provided onthe cloud 2407. Diabetic patients, for example, would probably bewilling to pay a periodic fee for monitoring their glucose levels,storing the history of the glucose levels, and having alarm warningswhen the glucose level falls out of range. Similarly, patients takingketone bodies supplement for treatment of disorders characterized byimpaired glucose metabolism (e.g., Alzheimer's, Parkinson's,Huntington's or ALS) may need to monitor their ketone bodies level.These patients would also probably be willing to pay a periodic fee forthe value-add services provided on the cloud 2407. Thus, by leveragingthe advances in wireless connectivity and the widespread use of handhelddevices such as smart phones that can wirelessly connect to the cloud,businesses can build a recurring cost business model even usingnon-invasive measurement devices.

In addition, it may be advantageous to use pattern matching algorithmsand other software and mathematical methods to identify the bloodconstituents of interest. In one embodiment, the spectrum may becorrelated with a library of known spectra to determine the overlapintegrals, and a threshold function may be used to quantify theconcentration of different constituents. This is just one way to performthe signal processing, and many other techniques, algorithms, andsoftware may be used and would fall within the scope of this disclosure.

Described herein are just some examples of the beneficial use ofnear-infrared or SWIR lasers for non-invasive monitoring of glucose,ketones, HbA1c and other blood constituents. However, many other medicalprocedures can use the near-infrared or SWIR light consistent with thisdisclosure and are intended to be covered by the disclosure.

In another specific embodiment, experiments have been performed forstand-off detection of solid targets with diffuse reflectionspectroscopy using a fiber-based super-continuum source (furtherdescribed herein). In particular, the diffuse reflection spectrum ofsolid samples such as explosives (TNT, RDX, PETN), fertilizers (ammoniumnitrate, urea), and paints (automotive and military grade) have beenmeasured at stand-off distances of 5 m. Although the measurements weredone at 5 m, calculations show that the distance could be anywhere froma few meters to over 150 m. These are specific samples that have beentested, but more generally other materials (particularly comprisinghydro-carbons) could also be tested and identified using similarmethods. The experimental set-up 2500 for thereflection-spectroscopy-based stand-off detection system is shown inFIG. 25, while details of the SC source 2501 are described in thisdisclosure in FIGS. 20,21, and 23. First, the diverging SC output iscollimated to a 1 cm diameter beam using a 25 mm focal length, 90degrees off-axis, gold coated, parabolic mirror 2502. To reduce theeffects of chromatic aberration, refractive optics are avoided in thesetup. All focusing and collimation is done using metallic mirrors thathave almost constant reflectivity and focal length over the entire SCoutput spectrum. The sample 2504 is kept at a distance of 5 m from thecollimating mirror 2502, which corresponds to a total round trip pathlength of 10 m before reaching the collection optics 2505. A 12 cmdiameter silver coated concave mirror 2505 with a 75 cm focal length iskept 20 cm to the side of the collimation mirror 2502. The mirror 2505is used to collect a fraction of the diffusely reflected light from thesample, and focus it into the input slit of a monochromator 2506. Thus,the beam is incident normally on the sample 2504, but detected at areflection angle of tan−1(0.2/5) or about 2.3 degrees. Appropriate longwavelength pass filters mounted in a motorized rotating filter wheel areplaced in the beam path before the input slit 2506 to avoid contributionfrom higher wavelength orders from the grating (300 grooves/mm, 2 μmblaze). The output slit width is set to 2 mm corresponding to a spectralresolution of 10.8 nm, and the light is detected by a 2 mm×2 mm liquidnitrogen cooled (77K) indium antimonide (InSb) detector 2507. Thedetected output is amplified using a trans-impedance pre-amplifier 2507with a gain of about 105V/A and connected to a lock-in amplifier 2508setup for high sensitivity detection. The chopper frequency is 400 Hz,and the lock-in time constant is set to 100 ms corresponding to a noisebandwidth of about 1 Hz. These are exemplary elements and parametervalues, but other or different optical elements may be used consistentwith this disclosure.

Process Analytical Technology (PAT)

One definition of process analytical technology, PAT, is “a system fordesigning, analyzing and controlling manufacturing through timelyevaluations (i.e., during processing) of significant quality andperformance attributes of raw and in-process materials and processes,with the goal of ensuring final product quality.” Near-infrared or SWIRspectroscopy may have applications in the PAT of the pharmaceuticalindustry by providing, for example, quantitative analysis of multiplecomponents in a sample and in pack quantification of drugs informulation, as well as quality of a drug and quality control of complexexcipients used in formulation. The PAT process may benefit fromnear-infrared or SWIR spectroscopy for some steps, such as: raw materialidentification, active pharmaceutical ingredient applications, drying,granulation, blend uniformity and content uniformity. Some of thestrengths of near-infrared or SWIR spectroscopy include: radiation hasgood penetration properties, and, thus, minimal sample preparation maybe required; measurement results may be obtained rapidly, andsimultaneous measurements may be obtained for several parameters;non-destructive methods with little or no chemical waste; and organicchemicals that comprise most pharmaceutical products have unique spectrain the near-infrared and SWIR ranges, for example.

One goal of the manufacturing process and PAT may be the concept of a“smart” manufacturing process, which may be a system or manufacturingoperation responding to analytical data generated in real-time. Such asystem may also have an in-built “artificial intelligence” as decisionsmay be made whether to continue a manufacturing operation. For example,with respect to the raw materials, integration of the qualitymeasurement into smart manufacturing processes could be used to improvemanufacturing operations by ensuring that the correct materials of theappropriate quality are used in the manufacture. Similarly, a smartblender would be under software control and would respond to thereal-time spectral data collected.

FIG. 26 illustrates what might be an eventual flow-chart 2600 of a smartmanufacturing process. The manufacturing process 2601 may have as inputthe process feed 2602 and result in a process output 2603. A processcontroller 2604 may at least partially control the manufacturing process2601, and the controller 2604 may receive inputs from the closed loopcontrol (process parameters) 2605 as well as the on-line monitoring ofprocess parameters 2606. The feedback loops in the process could refinethe manufacturing process 2601 and improve the quality of the processoutput 2603. These are particular embodiments of the use ofnear-infrared or SWIR spectroscopy in the PAT of the pharmaceuticalindustry, but other variations, combinations, and methods may also beused and are intended to be covered by this disclosure.

The discussion thus far has included use of near-infrared or SWIRspectroscopy in applications such as identification of counterfeitdrugs, detection of illicit drugs, and pharmaceutical process control.Although drugs and pharmaceuticals are one example, many other fieldsand applications may also benefit from the use of near infrared or SWIRspectroscopy, and these may also be implemented without departing fromthe scope of this disclosure. As just another example, near-infrared orSWIR spectroscopy may also be used as an analytic tool for food qualityand safety control. Applications in food safety and quality assessmentinclude contaminant detection, defect identification, constituentanalysis, and quality evaluation. The techniques described in thisdisclosure are particularly valuable when non-destructive testing isdesired at stand-off or remote distances.

Cancer and Other Diagnostics

Breast cancer is considered to be the most common cancer among women inindustrialized countries. It is believed that early diagnosis andconsequent therapy could significantly reduce mortality. Mammography isconsidered the gold standard among imaging techniques in diagnosingbreast pathologies. However, the use of ionizing radiation inmammography may have adverse effects and lead to other complications.Moreover, screening x-ray mammography may be limited by false positivesand negatives, leading to unnecessary physical and psychologicalmorbidity. Although breast cancer is one of the focuses of thisdisclosure, the same techniques may also be applied to other cancertypes, including, for example, skin, prostate, brain, pancreatic, andcolorectal cancer.

Diagnostic methods for assessment and therapy follow-up of breast cancerinclude mammography, ultrasound, and magnetic resonance imaging. Themost effective screening technique at this time is x-ray mammography,with an overall sensitivity for breast cancer detection around 75%,which is even further reduced in women with dense breasts to around 62%.Moreover, x-ray mammography has a 22% false positive rate in women under50, and the method cannot accurately distinguish between benign andmalignant tumors. Magnetic resonance imaging and ultrasound aresometimes used to augment x-ray mammography, but they have limitationssuch as high cost, low throughput, limited specificity and lowsensitivity. Thus, there is a continued need to detect cancers earlierfor treatment, missed by mammography, and to add specificity to theprocedures.

Optical breast imaging may be an attractive technique for breast cancerto screen early, augment with mammography, or use in follow-ontreatments. Also, optical breast imaging may be performed by intrinsictissue contrast alone (e.g., hemoglobin, water, collagen, and lipidcontent), or with the use of exogenous fluorescent probes that targetspecific molecules. For example, near-infrared (NIR) light may be usedto assess optical properties, where the absorption and scattering by thetissue components may change with carcinoma. For most of the studiesconducted to date, NIR light in the wavelength range of 600-1000 nm hasbeen used for sufficient tissue penetration; these wavelengths havepermitted imaging up to several centimeters deep in soft tissue. Opticalbreast imaging using fluorescent contrast agents may improve lesioncontrast and may potentially permit detection of changes in breasttissue earlier. In one embodiment, the fluorescent probes may eitherbind specifically to certain targets associated with cancer or maynon-specifically accumulate at the tumor site.

Optical methods of imaging and spectroscopy can be non-invasive usingnon-ionizing electromagnetic radiation, and these techniques could beexploited for screening of wide populations and for therapy monitoring.“Optical mammography” may be a diffuse optical imaging technique thataims at detecting breast cancer, characterizing its physiological andpathological state, and possibly monitoring the efficacy of thetherapeutic treatment. The main constituents of breast tissue may belipid, collagen, water, blood, and other structural proteins. Theseconstituents may exhibit marked and characteristic absorption featuresin the NIR wavelength range. Thus, diffuse optical imaging andspectroscopy in the NIR may be helpful for diagnosing and monitoringbreast cancer. Another advantage of such imaging is that opticalinstruments tend to be portable and more cost effective as compared toother instrumentation that is conventionally used for medical diagnosis.This can be particularly true, if the mature technologies fortelecommunications and fiber optics are exploited.

Spectroscopy using NIR or short-wave infrared (SWIR) light may bebeneficial, because most tissue has organic compounds that have overtoneor combination absorption bands in this wavelength range (e.g., betweenapproximately 0.8-2.5 microns). In one embodiment, a NIR or SWIRsuper-continuum (SC) laser that is an all-fiber integrated source may beused as the light source for diagnosing cancerous tissue. Exemplaryfiber-based super-continuum sources may emit light in the NIR or SWIRbetween approximately 1.4-1.8 microns, 2-2.5 microns, 1.4-2.4 microns,1-1.8 microns, or any number of other bands. In particular embodiments,the detection system may be one or more photo-detectors, a dispersivespectrometer, a Fourier transform infrared spectrometer, or ahyper-spectral imaging detector or camera. In addition, reflection ordiffuse reflection light spectroscopy may be implemented using the SWIRlight source, where the spectral reflectance can be the ratio ofreflected energy to incident energy as a function of wavelength.

For breast cancer, experiments have shown that with growing cancer thecollagen content increases while the lipid content decreases. Therefore,early breast cancer detection may involve the monitoring of absorptionor scattering features from collagen and lipids. In addition, NIRspectroscopy may be used to determine the concentrations of hemoglobin,water, as well as oxygen saturation of hemoglobin and optical scatteringproperties in normal and cancerous breast tissue. For optical imaging tobe effective, it may also be desirable to select the wavelength rangethat leads to relatively high penetration depths into the tissue. In oneembodiment, it may be advantageous to use optical wavelengths in therange of about 1000-1400 nm. In another embodiment, it may beadvantageous to use optical wavelengths in the range of about 1600-1800nm. Higher optical power densities may be used to increase thesignal-to-noise ratio of the detected light through the diffusescattering tissue, and surface cooling or focused light may bebeneficial for preventing pain or damage to the skin and outer layersurrounding the breast tissue. Since optical energy may be non-ionizing,different exposure times may be used without danger or harmfulradiation.

To perform non-invasive optical mammography, one desired attribute isthat the light may penetrate as far as possible into the breast tissue.In diffuse reflection spectroscopy, a broadband light spectrum may beemitted into the tissue, and the spectrum of the reflected ortransmitted light may depend on the absorption and scatteringinteractions within the target tissue. Since breast tissue hassignificant water and hemoglobin content, it is valuable to examine thewavelength range over which deep penetration of light is possible. FIG.27 illustrates the optical absorption 2700 of pure water (dotted line)2701, hemoglobin without oxygen (thinner solid line) 2702, andhemoglobin saturated with oxygen (thicker solid line) 2703. It can benoted that above about 1100 nm, the absorption of hemoglobin is almostthe same as water absorption. The penetration depth may be proportionalto the inverse of the optical absorption. Therefore, the highestpenetration depth will be at the absorption valley, approximately in thewavelength range between about 900 nm and about 1300 nm. Although not aslow in absorption compared to the first window, another absorptionvalley lies between about 1600 nm and 1800 nm. Thus, non-invasiveimaging preferably should use wavelengths that fall in one of these twoabsorption valleys.

FIG. 28 shows examples of various absorption bands of chemical species2800 in the wavelength range between about 1200 nm and 2200 nm. Althoughthe fundamental absorptions usually lie in the mid-infrared (e.g.,wavelengths longer than about 3 microns), there are many absorptionlines in the NIR corresponding to the second overtone region 2801between about 1000 nm and 1700 nm, the first overtone region 2802between about 1500 nm and 2050 nm, and the combination band region 2803between about 1900 nm and 2300 nm. As an example, hydrocarbon bondscommon in many biological substances have their fundamental absorptionin the mid-IR near 3300-3600 nm, but they also have many combinationband lines between 2000-2500 nm, and other lines at shorter wavelengthscorresponding to the first and second overtones. Fortunately, there arespectral features of FIG. 28 that overlap with the absorption valleys inFIG. 27. These are likely to be the wavelengths of interest forspectroscopic analysis of cancerous regions.

In women, the breasts (FIG. 29) 2900 overlay the pectoralis majormuscles 2902 and cover much of the chest area and the chest walls 2901.The breast is an apocrine gland that produces milk to feed an infantchild; the nipple 2904 of the breast is surrounded by an areola 2905,which has many sebaceous glands. The basic units of the breast are theterminal duct lobules 2903, which produce the fatty breast milk. Theygive the breast its function as a mammary gland. The lobules 2903 feedthrough the milk ducts 2906, and in turn these ducts drain to the nipple2904. The superficial tissue layer (superficial fascia) may be separatedfrom the skin 2908 by about 0.5-2.5 cm of adipose of fatty tissue 2907.

Breast cancer is a type of cancer originating from breast tissue, mostcommonly from the inner lining of milk ducts 2906, the lobules 2903 thatsupply the ducts with milk, and/or the connective tissue between thelobules. Cancers originating from ducts 2906 are known as ductalcarcinomas, while those originating from lobules 2903 or theirconnective tissue are known as lobular carcinomas. While theoverwhelming majority of human cases occur in women, male breast cancermay also occur.

Several particular embodiments of imaging systems 3000, 3050 foroptically scanning a breast are illustrated in FIG. 30. In theseparticular embodiments, the patient 3001, 3051 may lie in a proneposition with her breasts inside a box 3002, 3052 with probably atransparent window on the detector side. A compression plate 3003, 3053may hold the breast in place against the viewing window by mildlycompressing the breast to a thickness between about 5.5 and 7.5 cm. Thebox 3002, 3052 may then be filled with a matching fluid with opticalproperties similar to human breast. In one instance, the matching fluidmay comprise water, india ink for absorption, and a fat emulsion forscattering. The embodiments in FIG. 30 may also have one or moredetectors 3004, 3055, one or more light sources 3004, 3054, variouselectronics, and even an imaging system based on charge coupled devices3005. As illustrated in FIG. 30, the light sources 3004, 3054 anddetectors 3004, 3055 may be coupled to the box 3002, 3052 through one ormore fibers 3006, 3056. Also, the imaging may be in reflection mode (topof FIG. 30), transmission mode (bottom of FIG. 30), or some combination.

Beyond the geometry and apparatus of FIG. 30, the optical imaging systemmay use one or more of three different illumination methods: continuouswave, time-domain photon migration, and frequency-domain photonmigration. In one embodiment, continuous-wave systems emit light atapproximately constant intensity or modulated at low frequencies, suchas 0.1-100 kHz. In another embodiment, the time-domain photon migrationtechnique uses relatively short, such as 50-400 psec, light pulses toassess the temporal distribution of photons. Since scattering mayincrease the times of flight spent by photons migrating in tissues, thephotons that arrive earliest at the detector probably encountered thefewest scattering events. In yet another embodiment, thefrequency-domain photon migration devices modulate the amplitude of thelight that may be continuously transmitted at relatively highfrequencies, such as 10 MHz to 1 GHz. For example, by measuring thephase shift and amplitude decay of photons as compared to a referencesignal, information may be acquired on the optical properties of tissue,and scattering and absorption may be distinguished. Beyond these threemethods, other techniques or combinations of these methods may be used,and these other methods are also intended to fall within the scope ofthis disclosure.

Although particular embodiments of imaging architectures are illustratedin FIG. 30, other system architectures may also be used and are alsointended to be covered by this disclosure. For example, in oneembodiment several couples of optical fibers for light delivery andcollection may be arranged along one or more rings placed at differentdistances from the nipple 2904. In an alternate embodiment a “cap” withfiber leads for light sources and detectors may be used that fits overthe breast. In yet another embodiment, imaging optics and light sourcesand detectors may surround the nipple 2904 and areola 2905 regions ofthe breast. As yet another alternative, a minimally invasive proceduremay involve inserting needles with fiber enclosure (to light sources anddetectors or receivers) into the breast, so as to probe regions such asthe lobules 2903 and connective tissue. Both non-invasive and minimallyinvasive optical imaging methods are intended to be covered by thisdisclosure.

Optical Wavelength Ranges for Cancer Detection

Many of the diffuse optical tomography studies previously conducted haverelied on using NIR in the wavelength range of about 600-1000 nm, wherelight absorption at these wavelengths may be minimal, allowing forsufficient tissue penetration (up to 15 cm). In these wavelength ranges,it has been claimed that concentrations of oxy- and deoxy-hemoglobin,water, and lipids can be determined. For example, FIG. 31 shows thenormalized absorption spectra 3100 of main tissue absorbers in the NIRbetween about 600 nm and 1100 nm: deoxy-hemoglobin, Hb, 3101,oxy-hemoglobin, HbO2, 3102, water 3103, lipids 3104 and collagen 3105.It is speculated that in a malignant tumor, hemoglobin concentration maybe directly related to angiogenesis, one of the main factors requiredfor tumor growth and metastases. Moreover, the proportions of oxy- anddeoxy-hemoglobin in a tumor may change due to its metabolism. Thus, bymeasuring concentrations of the breast components, discrimination ofbenign and malignant tumors may be possible with diffuse opticalimaging. Experiment evidence suggests that cancerous tissue isassociated with higher hemoglobin and water concentrations, and a lowerlipid concentration with respect to normal breast tissue.

Based on FIG. 27 and the dynamics of carcinoma, it may be advantageousto perform spectroscopy in longer wavelengths, such as windows between1000-1400 nm or 1600-1800 nm. For example, looking at the absorptioncurves 2700 in FIG. 27, the absorption between approximately 1000-1300nm may be comparable to the 600-1000 nm window described in FIG. 31.However, the loss through the soft tissue medium (penetration depth maybe inversely related to the loss) will be due to absorption andscattering. In fact, the scattering properties of tissue may alsocontain valuable information for lesion diagnosis. Since the scatteringis inversely proportional to some power of wavelength (for example, insome tissue scattering is inversely proportional to the wavelengthcubed), the scattering contribution to the loss may decrease at longerwavelengths. Moreover, these longer wavelength windows may containdiagnostic information on content of collagen and adipose, both of whichmay be significant indicators for breast cancer.

Breast cancer spectroscopy may benefit from the use of wavelengthslonger than about 1000 nm for a number of reasons. As one example, themain absorbers in soft tissues of the visible spectrum of light may beoxy- and deoxygenated hemoglobin and beta-carotene. On the other hand,primary absorbers in the near-infrared spectrum of light may be water,adipose tissue and collagen. Particularly adipose and collagen contentmay be valuable for early detection of cancers. In one embodiment,increased levels of collagen in breast malignancies are thought to bedue to increased vascularity of the tumors. Collagen type I may be animportant component of artery walls. FIG. 32 illustrates the normalizedabsorption coefficient 3200 in the wavelength range between about 500 nmand 1600 nm for Hb 3201, HbO2 3202, beta-carotene 3203, water 3204,lipid 3205 and collagen 3206.

Collagen and Adipose Signatures in Near-IR

Examining the collagen content may be a valuable indicator for breastcancer detection. Collagen is one of the important extracellular matrixproteins, and fibrillar collagens help to determine stromalarchitecture. In turn, changes in the stromal architecture andcomposition are one of the aspects of both benign and malignantpathologies, and, therefore, may play an initial role in breastcarcinogenesis. For example, collagen seems to be related to cancerdevelopment, because high mammographic density may be recognized as arisk factor for breast cancer. Moreover, collagen type in high-riskdense breasts may appear to be different from collagen in low-densitybreasts.

Experimental data also shows that malignant mammary gland tissues ofanimals and humans show a decrease in lipids when compared to normaltissues. The reduced amounts of lipids in the cancerous sites may becaused by a high metabolic demand of lipids in the malignant tumors. Forexample, due to the rapid proliferation of cancerous cells, there may bereduced lipid content in cancerous tissues. Thus, in addition tocollagen, another valuable marker for breast cancer may be the lipidspectral features. It may also be possible to combine the markers fromoxy- and deoxygenated hemoglobin and water with lipid and collagen linesto improve the diagnostics and/or therapeutics of optical imaging and/ortreatment for breast and other types of cancer. Although specificexamples of tissue constituents are discussed, other tissue constituentsand related markers may also be associated with breast cancer and othercancers, and these other constituents are also intended to be covered bythis disclosure.

As an example of the types of spectral signatures that may exist, invivo investigations of progressive changes in rat mammary gland tumorswere conducted using near-infrared spectroscopy with a Fourier-transforminfrared spectrometer. In one embodiment, FIG. 33 shows the typicalspectra of the cancerous site of the treated rat and the correspondingsite of the normal rat. FIG. 33A shows the logarithm of the inverse ofreflection spectra 3300, while FIG. 33B shows their second derivativespectra 3350. The curves 3301, 3351 correspond to the spectra of thecancerous sites, while 3302, 3352 correspond to the spectra of thenormal sites. Since the second derivative techniques may be useful inthe analyses of NIR spectra to minimize baseline offsets and to resolveoverlapping absorption without compromising signal-to-noise, FIG. 33Bmay be used for interpretation of the spectral changes.

In FIG. 33B identification may be made of several of the spectralfeatures. In particular, there are DNA bands near 1471 nm and 1911 nm,while there are water bands near 967 nm, 1154 nm, 1402 nm, and 1888 nm.Moreover, there are lipid bands near 1209 nm, 1721 nm and 1764 nm, andthere are protein bands near 2055 nm, 2172 nm and 2347 nm. The NIRspectra of FIG. 33 show that the DNA and water contents in the canceroustissue may be higher than those in normal tissues. On the other hand,the lipid content in the cancerous tissue may be less than the lipidcontent in normal tissues. With protein contents, however, littledifference may be found between the normal and cancerous tissue.

These experiments on rats with breast cancer were also used to observethe temporal progression of the cancer. In this embodiment, as thecancer grew, the lipid band intensity decreased, and this band alsoshifted to higher wavelengths, and collagen peaks appeared in thetissues. In FIG. 34, the second derivative spectral changes 3400 wereinvestigated in the 1600 nm to 1800 nm wavelength range over severalweeks. An early cancer was detected in the 5th week, and then it grewrapidly from the 6th 3401 to the 7th 3402 week. The cancer's temporalprogression through the 8th 3403, 9th 3404, 10th 3405 and 11th 3406 weekare shown in the various curves in FIG. 34. With the cancer growth, theintensity of the lipid band in the vicinity of 1721 nm decreased, andthis band shifted to higher wavelengths by 7 nm at the 11th week 3406compared to the wavelength band at the 5th week. The higher wavelengthshift may indicate that an order parameter of the lipids increases withprogressive cancer growth.

Moreover, in the data of FIG. 34 is seen that a new peak appeared as thecancer grew around 1690 nm, which may be assigned to be a collagenabsorption by comparison with the absorptions of standard collagen(c.f., FIG. 37). The higher wavelength shift may be attributable to theformation of elastic fibers in the lipid layer with collagen induced inthe cancer tissues, thus leading to an increased order parameter of thelipids. Thus, it can be seen that significant information about breastcancer tissue compared with normal tissue may be obtained byspectroscopy at the longer wavelengths in the near-infrared.

The second derivative spectra may also be insightful for observing andmonitoring changes in tissue as well as characterizing tissue in thenear-infrared wavelength range. As an example, FIG. 35 illustrates thesecond derivative spectra 3500 for cholesterol (similar to oneembodiment of lipids) 3501, collagen 3502, and elastin 3503. The leftcurve 3525 shows the second derivative data over the wavelength range ofabout 1150 nm to 1300 nm, while the right curve 3550 shows the secondderivative data over the wavelength range of about 1600 nm to 1850 nm.These wavelengths show numerous features for cholesterol/lipid 3501,collagen 3502, and elastin 3503, which again emphasizes the added valueof using wavelengths longer than about 1000 nm for cancer diagnostics.

To further illustrate the value of using longer wavelengths in the NIRor SWIR for observing changes in breast cancer and other cancer markers,the spectra of in water, lipids/adipose and collagen of differentvarieties may be studied. As one embodiment, the absorption coefficients1000 are shown in FIG. 36 as a function of wavelength between about 1000nm and 2600 nm. FIG. 36 overlaps the absorption coefficient for water3601, adipose 3602 (forms of adipose include fatty tissue and acids,lipids, and cholesterol), and collagen type I 3603. One may note thatparticular absorption peaks for adipose 3602 and collagen type I 3603align at wavelengths near 1210 nm 3604 and 1720 nm 3605, which alsocorrespond to local minima in water absorption.

Moreover, the NIR spectra for collagen also depend on the type ofcollagen. As an example, FIG. 37 illustrates the absorbance 3700 forfour types of collagen: collagen I 3701, collagen II 3702, collagen III3703, and collagen IV 3704. Collagen I, for instance, may be a majorconstituent of stroma. Also, collagen I and collagen III may be theprincipal collagens of the aorta. Since the spectra of the fourcollagens are distinctive, multicomponent analysis of collagens maypossibly be used to distinguish the type of collagen involved.

The experimental results discussed thus far indicate that breast cancerdetection may benefit from spectroscopy in the NIR and SWIR,particularly wavelengths between approximately 1000-1400 nm and1600-1800 nm. These are wavelength windows that may have deeppenetration into soft tissue, while still falling within lowerabsorption valleys of water. Moreover, the longer wavelengths lead toless scattering in tissue and water, again permitting deeper penetrationof the light. In the NIR and SWIR wavelength range, the spectra ofstandard samples of cholesterol, protein, collagen, elastin and DNA weremeasured to obtain information on their characteristic bands in thespectra of mammary gland tissues. Absorption peaks in the standardsamples occur at the following exemplary wavelengths:

Comparing these absorption features with the data in FIGS. 32-37 showsthat there are absorption features or signatures in the secondderivatives that can be used to monitor changes in, for example,collagen and lipids. By using broadband light and performingspectroscopy in at least some part of the wavelength windows betweenabout 1000-1400 nm and/or 1600-1800 nm, the collagen and lipid changes,or other constituent changes, may be monitored. In one embodiment, forbreast cancer the decrease in lipid content, increase in collagencontent, and possible shift in collagen peaks may be observed byperforming broadband light spectroscopy and comparing normal regions tocancerous regions as well as the absorption strength as a function ofwavelength. The spectroscopy may be in transmission, reflection, diffusereflection, diffuse optical tomography, or some combination. Also, thisspectroscopy may be augmented by fluorescence data, if particular tagsor markers are added. Beyond looking at the absorbance, the dataprocessing may involve also observing the first, second, or higher orderderivatives.

Broadband spectroscopy is one example of the optical data that can becollected to study breast cancer and other types of cancer. However,other types of spectral analysis may also be performed to compare thecollagen and lipid features between different wavelengths and differenttissue regions (e.g., comparing normal regions to cancerous regions),and these methods also fall within the scope of this disclosure. Forexample, in one embodiment just a few discrete wavelengths may bemonitored to see changes in lipid and collagen contents. In a particularembodiment, wavelengths near 1200 nm may be monitored in the secondderivative data of FIG. 35 to measure the cholesterol/lipid peak below1200 nm in 3501 versus the collagen peak above 1200 nm in 3502. In yetanother embodiment, the absorption features in FIG. 32 may be reliedupon to monitor the lipid content 3205 by measuring near 1200 nm and thecollagen content 3206 by measuring near 1300 nm. Although theseembodiments use only two wavelengths, any number of wavelengths may beused and are intended to be covered by this disclosure.

Thus, a breast cancer monitoring system, or a system to monitordifferent types of cancers, may comprise broadband light sources anddetectors to permit spectroscopy in transmission, reflection, diffuseoptical tomography, or some combination. In one particular embodiment,high signal-to-noise ratio may be achieved using a fiber-basedsuper-continuum light source (described further herein). Other lightsources may also be used, including a plurality of laser diodes,super-luminescent laser diodes, or fiber lasers.

Wavelength ranges that may be advantageous for cancer detection includethe NIR and SWIR windows (or some part of these windows) between about1000-1400 nm and 1600-1800 nm. These longer wavelengths fall withinlocal minima of water absorption, and the scattering loss decreases withincreasing wavelength. Thus, these wavelength windows may permitrelatively high penetration depths. Moreover, these wavelength rangescontain information on the overtone and combination bands for variouschemical bonds of interest, such as hydrocarbons.

These longer wavelength ranges may also permit monitoring levels andchanges in levels of important cancer tissue constituents, such aslipids and collagen. Breast cancer tissue may be characterized bydecreases in lipid content and increases in collagen content, possiblywith a shift in the collagen peak wavelengths. The changes in collagenand lipids may also be augmented by monitoring the levels of oxy- anddeoxy-hemoglobin and water, which are more traditionally monitoredbetween 600-1000 nm. Other optical techniques may also be used, such asfluorescent microscopy.

To permit higher signal-to-noise levels and higher penetration depths,higher intensity or brightness of light sources may be used. With thehigher intensities and brightness, there may be a higher risk of pain orskin damage. At least some of these risks may be mitigated by usingsurface cooling and focused infrared light, as further described herein.

Laser Experiments: Penetration Depth, Focusing, Skin Cooling

Some preliminary experiments show the feasibility of using focusedinfrared light for non-invasive procedures, or other procedures whererelatively shallow vessels below the skin are to be thermally coagulatedor occluded with minimum damage to the skin upper layers. In oneembodiment, the penetration depth and optically induced thermal damagehas been studied in chicken breast samples. Chicken breast may be areasonable optical model for smooth muscle tissue, comprising water,collagen and proteins. Commercially available chicken breast sampleswere kept in a warm bath (˜32 degree Celsius) for about an hour, andthen about half an hour at room temperature in preparation for themeasurements.

An exemplary set-up 3800 for testing chicken breast samples usingcollimated light is illustrated in FIG. 38. The laser light 3801 near980 nm, 1210 nm, or 1700 nm may be provided from one or more laserdiodes or fiber lasers, as described further below. In this instance,laser diodes were used, which comprise a plurality of laser diodeemitters that are combined using one or more multiplexers (particularlyspatial multiplexers), and then the combined beam is coupled into amulti-mode fiber (typically 100 microns to 400 microns in diameter). Theoutput from the laser diode fiber was then collimated using one or morelenses 3802. The resulting beam 3803 was approximately round with a beamdiameter of about 3 mm. The beam diameter was verified by blademeasurements (i.e., translating a blade across the beam). Also, thetime-averaged power was measured in the nearly collimated section afterthe lens using a large power meter. The chicken breast samples 3806 weremounted in a sample holder 3805, and the sampler holder 3805 was mountedin turn on a translation stage 3804 with a linear motor that could moveperpendicular to the incoming laser beam. Although particular details ofthe experiment are described, other elements may be added or eliminated,and these alternate embodiments are also intended to be covered by thisdisclosure.

For these particular experiments, the measured depth of damage (inmillimeters) versus the incident laser power (in Watts) is shown 3900 inFIG. 39. In this embodiment, laser diodes were used at wavelengths near980 nm, 1210 nm and 1700 nm. The curve 3901 corresponds to about 980 nm,the curve 3902 corresponds to about 1210 nm, and the curve 3903corresponds to about 1700 nm. It may be noted that there is a thresholdpower, above which the damage depth increases relatively rapidly. Forexample, the threshold power for wavelengths around 980 nm may be about8 W, the threshold power for wavelengths around 1210 nm may be 3 W, andthe threshold power for wavelengths around 1700 nm may be about 1 W. Thethreshold powers may be different at the different wavelengths becauseof the difference in water absorption (e.g., 3601 in FIG. 36). Part ofthe difference in threshold powers may also arise from the absorption ofproteins such as collagen (e.g., 3603 in FIG. 36). After a certain powerlevel, the damage depth appears to saturate: i.e., the slope flattensout as a function of increasing pump power.

In one embodiment, if the penetration depth is defined as the depthwhere damage begins to approximately saturate, then for wavelengths ofabout 980 nm 3901 the penetration depth 3906 may be defined asapproximately 4 mm, for wavelengths of about 1210 nm 3902 thepenetration depth 3905 may be defined as approximately 3 mm, and forwavelengths of about 1700 nm 3903 the penetration depth 3904 may bedefined as approximately 2 mm. These are only approximate values, andother values and criteria may be used to define the penetration depth.It may also be noted that the level of damage at the highest powerpoints differs at the different wavelengths. For example, at the highestpower point of 3903 near 1700 nm, much more damage is observed, showingevidence of even boiling and cavitation. This may be due to the higherabsorption level near 1700 nm (e.g., 3601 in FIG. 36). On the otherhand, at the highest power point 3901 near 980 nm, the damage is not ascatastrophic, but the spot size appears larger. The larger spot size maybe due to the increased scattering at the shorter wavelengths (e.g.,3601 in FIG. 36). Based on data 3900 such as in FIG. 39, it may bepossible to select the particular wavelength for the laser beam to beused in the non-invasive procedure.

Even near wavelengths such as described in FIG. 39, the particularwavelength selected may be more specifically defined based on the targettissue of interest. In one particular embodiment, the vessel lumen maybe modeled as water, and for this example assume that wavelengths in thevicinity of 980 nm are being selected to create thermal coagulation orocclusion. FIG. 40 shows the optical absorption or density as a functionof wavelength 4000 between approximately 700 nm and 1300 nm. Curves areshown for the water absorption 4001, hemoglobin Hb absorption 4002, andoxygenated hemoglobin HbO2 4003. In this example, two particularwavelengths are compared: 980 nm 4004 and 1075 nm 4005. For instance,980 nm may be generated using one or more laser diodes, while 1075 nmmay be generated using an ytterbium-doped fiber laser. If maximizing thepenetration depth is the significant problem, then 1075 nm 4005 may bepreferred, since it falls near a local minimum in water 4001, hemoglobin4002, and oxygenated hemoglobin 4003 absorption. On the other hand, ifthe penetration depth at 980 nm 4004 is adequate and the problem is togenerate heat through water absorption, then 980 nm 4004 may be apreferred wavelength for the light source because of the higher waterabsorption. This wavelength range is only meant to be exemplary, butother wavelength ranges and particular criteria for selecting thewavelength may be used and are intended to be covered by thisdisclosure.

In another embodiment, focused infrared light has been used to preservethe top layer of a tissue while damaging nerves at a deeper level. Forinstance, FIG. 41 illustrates the set-up 4100 used for the focusedinfrared experiments. In this embodiment, a lens 4101 is used to focusthe light. Although a single lens is shown, either multiple lenses, GRIN(gradient index) lenses, curved mirrors, or a combination of lenses andmirrors may be used. In this particular example, the tissue 4104 isplaced between two microscope slides 4102 and 4103 for in vitroexperiments. The tissue 4104 is renal artery wall either from porcine orbovine animals (about 1.2 mm thick sample)—i.e., this is the arteryleading to the kidneys, and it is the artery where typically renaldenervation may be performed to treat hypertension. For this example,the minimum beam waist 4105 falls behind the tissue, and the intensitycontrast from the front of the tissue (closest to the lens) to the backof the tissue (furthest from the lens) is about 4:1. These areparticular ranges used for this experiment, but other values andlocations of minimum beam waist may also be used and intended to becovered by this disclosure.

For a particular embodiment, histology of the renal artery is shown inFIG. 42A for no laser exposure 4200 and shown in FIG. 42B with focusedinfrared laser exposure 4250. In this experiment, the beam diameterincident on the lens was about 4 mm, and the distance from the edge ofthe flat side of lens to the minimum beam waist was about 3.75 mm. Thebeam diameter on the front side of the renal artery (i.e., theendothelium side) was about 1.6 mm, and the beam diameter on the backside of the renal artery was about 0.8 mm. In FIG. 42A with no laserexposure, the layers of the artery wall may be identified: top layer ofendothelium 4201 that is about 0.05 mm thick, the media comprisingsmooth muscle cells or tissue 4202 that is about 0.75 mm thick, and theadventitia 4203 comprising some of the renal nerves 4204 that is about1.1 mm thick. These are particular values for this experiment, and otherlayers and thicknesses may also be used and are intended to be coveredby this disclosure.

The histology with focused infrared light exposure 4250 is illustratedin FIG. 42B. The laser light used is near 1708 nm from a cascaded Ramanoscillator (described in greater detail herein), and the power incidenton the tissue is about 0.8 W and the beam is scanned across the tissueat a rate of approximately 0.4 mm/sec. The various layers are stillobservable: the endothelium 4251, the media 4252, and the adventitia4253. With this type of histology, the non-damaged regions remain darker(similar to FIG. 42A), while the laser induced damaged regions turnlighter in color. In this example, the endothelium 4251 and top layer ofthe media 4252 remain undamaged—i.e., the top approximately 0.5 mm isthe undamaged region 4256. The laser damaged region 4257 extends forabout 1 mm, and it includes the bottom layer of the media 4252 and muchof the adventitia 4253. The renal nerves 4254 that fall within thedamage region 4257 are also damaged (i.e., lighter colored). On theother hand, the renal nerves beyond this depth, such as 4255, may remainundamaged.

Thus, by using focused infrared light near 1708 nm in this example, thetop approximately 0.5 mm of the renal artery is spared from laserdamage. It should be noted that when the same experiment is conductedwith a collimated laser beam, the entire approximately 1.5 mm is damaged(i.e., including regions 4256 and 4257). Therefore, the cone of lightwith the lower intensity at the top and the higher intensity toward thebottom may, in fact, help preserve the top layer from damage. Thereshould be a Beer's Law attenuation of the light intensity as the lightpropagates into the tissue. For example, the light intensity shouldreduce exponentially at a rate determined by the absorption coefficient.In these experiments it appears that the focused light is able toovercome the Beer's law attenuation and still provide contrast inintensity between the front and back surfaces.

In another embodiment, experiments have also been conducted ondermatology samples with surface cooling, and surface cooling is shownto preserve the top layer of the skin during laser exposure. In thisparticular example, the experimental set-up 4300 is illustrated in FIG.43. The skin sample 4304, or more generally sample under test, is placedin a sample holder 4303. The sample holder 4303 has a cooling side 4301and a heating side 4302. The heating side 4302 comprises a heater 4305,which may be adjusted to operate around 37 degrees Celsius—i.e., closeto body temperature. The cooling side 4301 is coupled to an ice-waterbath 4307 (around 2 degrees Celsius) and a warm-water bath 4306 (around37 degrees Celsius) through a switching valve 4308. The entire sampleholder 4303 is mounted on a linear motor 4309, so the sample can bemoved perpendicular 4310 to the incoming light beam.

In this embodiment, the light is incident on the sample 4304 through asapphire window 4311. The sapphire material 4311 is selected because itis transparent to the infrared wavelengths, while also being a goodthermal conductor. Thus, the top layer of the sample 4304 may be cooledby being approximately in contact with the sapphire window 4311. Thelaser light 4312 used is near 1708 nm from a cascaded Raman oscillator(described in greater detail herein), and one or more collimating lenses4313 are used to create a beam with a diameter 4314 of approximately 2mm. This is one particular embodiment of the sample surface coolingarrangement, but other apparatuses and methods may be used and areintended to be covered by this disclosure.

Experimental results obtained using the set-up of FIG. 43 are includedin FIG. 44. In this example, FIG. 44 shows the MTT histochemistry ofhuman skin 4400 treated with ˜1708 nm laser (5 seconds pre-cool; 2 mmdiameter spot exposure for 3 seconds) at 725 mW (A 4401, B 4402)corresponding to about 70 J/cm2 average fluence, and 830 mW (C 4403, D4404) corresponding to about 80 J/cm2 average fluence. The images inFIG. 44 show that the application of a cold window was effective inprotecting the epidermis 4405 (darker top layer) and the topapproximately 0.4 or 0.5 mm of the dermis 4406. As before, the darkerregions of the histology correspond to undamaged regions, while thelighter regions correspond to damaged regions. In contrast, when nosurface cooling is applied, then thermal damage to the dermis occurs inthe epidermis and dermis where the laser exposure occurs, and thethermal damage extends to about 1.3 or 1.4 mm or more from the skinsurface. Thus, surface cooling applied to the skin may help to reduce oreliminate damage to the top layer of the skin under laser exposure.

In summary, experiments verify that infrared light, such as near 980 nm,1210 nm, or 1700 nm, may achieve penetration depths betweenapproximately 2 mm to 4 mm or more. The top layer of skin or tissue maybe spared damage under laser exposure by focusing the light beyond thetop layer, applying surface cooling, or some combination of the two.These are particular experimental results, but other wavelengths,methods and apparatuses may be used for achieving the penetration andminimizing damage to the top layer and are intended to be covered bythis disclosure. In an alternate embodiment, it may be beneficial to usewavelengths near 1310 nm if the absorption from skin constituents (FIG.36), such as collagen 3603, adipose 3602 and elastin 3604, are to beminimized. The water absorption 3601 near 1310 nm may still permit apenetration depth of approximately 1 cm, or perhaps less. In yet anotherembodiment, wavelengths near 1210 nm may be beneficial, if penetrationdepths on the order of 3 mm are adequate and less scattering loss (e.g.3601 in FIG. 36) is desired. Any of FIG. 27, 32, 35, 36, or 37 may beused to select these or other wavelengths to achieve the desiredpenetration depth and to also perhaps target particular tissue ofinterest, and these alternate embodiments are also intended to becovered by this disclosure.

Detection Systems

As discussed earlier, the active remote sensing system or hyper-spectralimaging system may be on an airborne platform, mounted on a vehicle, astationary transmission or reflection set-up, or even held by a humanfor a compact system. For such a system, there are fundamentally twohardware parts: the transmitter or light source and the detectionsystem. Between the two, perhaps in a transmission or reflectionsetting, may be the sample being tested or measured. Moreover, theoutput from the detection system may go to a computational system,comprising computers or other processing equipment. The output from thecomputational system may be displayed graphically as well as withnumerical tables and perhaps an identification of the materialcomposition. These are just some of the parts of the systems, but otherelements may be added or be eliminated, and these modifiedconfigurations are also intended to be covered by this disclosure.

By use of an active illuminator, a number of advantages may be achieved.First, the variations due to sunlight and time-of-day may be factoredout. The effects of the weather, such as clouds and rain, might also bereduced. Also, higher signal-to-noise ratios may be achieved. Forexample, one way to improve the signal-to-noise ratio would be to usemodulation and lock-in techniques. In one embodiment, the light sourcemay be modulated, and then the detection system would be synchronizedwith the light source. In a particular embodiment, the techniques fromlock-in detection may be used, where narrow band filtering around themodulation frequency may be used to reject noise outside the modulationfrequency. In an alternate embodiment, change detection schemes may beused, where the detection system captures the signal with the lightsource on and with the light source off. Again, for this system thelight source may be modulated. Then, the signal with and without thelight source is differenced. This may enable the sun light changes to besubtracted out. In addition, change detection may help to identifyobjects that change in the field of view. Using a lock-in type technique(e.g., detecting at the same frequency as the pulsed light source andalso possibly phase locked to the same signal), the detection system maybe able to reject background or spurious signals and increase thesignal-to-noise ratio of the measurement. In the following someexemplary detection systems are described.

In one embodiment, a SWIR camera or infrared camera system may be usedto capture the images. The camera may include one or more lenses on theinput, which may be adjustable. The focal plane assemblies may be madefrom mercury cadmium telluride material (HgCdTe), and the detectors mayalso include thermo-electric coolers. Alternately, the image sensors maybe made from indium gallium arsenide (InGaAs), and CMOS transistors maybe connected to each pixel of the InGaAs photodiode array. The cameramay interface wirelessly or with a cable (e.g., USB, Ethernet cable, orfiber optics cable) to a computer or tablet or smart phone, where theimages may be captured and processed. These are a few examples ofinfrared cameras, but other SWIR or infrared cameras may be used and areintended to be covered by this disclosure.

In one example of multi-beam detection systems, a dual-beam set-up 4500such as in FIG. 45 may be used to subtract out (or at least minimize theadverse effects of) light source fluctuations. In one embodiment, theoutput from an SC source 4501 may be collimated using a calcium fluoride(CaF2) lens 4502 and then focused into the entrance slit of themonochromator 4503. At the exit slit, light at the selected wavelengthis collimated again and may be passed through a polarizer 4504 beforebeing incident on a calcium fluoride beam splitter 4505. After passingthrough the beam splitter 4505, the light is split into a sample 4506and reference 4507 arm to enable ratiometric detection that may cancelout effects of intensity fluctuations in the SC source 4501. The lightin the sample arm 4506 passes through the sample of interest and is thenfocused onto a HgCdTe detector 4508 connected to a pre-amp. A chopper4502 and lock-in amplifier 4510 setup enable low noise detection of thesample arm signal. The light in the reference arm 4507 passes through anempty container (cuvette, gas cell etc.) of the same kind as used in thesample arm. A substantially identical detector 4509, pre-amp and lock-inamplifier 4510 is used for detection of the reference arm signal. Thesignal may then be analyzed using a computer system 4511. This is oneparticular example of a method to remove fluctuations from the lightsource, but other components may be added and other configurations maybe used, and these are also intended to be covered by this disclosure.

Although particular examples of detection systems have been described,combinations of these systems or other systems may also be used, andthese are also within the scope of this disclosure. As one example,environmental fluctuations (such as turbulence or winds) may lead tofluctuations in the beam for active remote sensing or hyper-spectralimaging. A configuration such as illustrated in the representativeembodiment of FIG. 45 may be able to remove the effect of environmentalfluctuations. Yet another technique may be to “wobble” the light beamafter the light source using a vibrating mirror. The motion may lead tothe beam moving enough to wash out spatial fluctuations within the beamwaist at the sample or detection system. If the vibrating mirror isscanned faster than the integration time of the detectors, then thespatial fluctuations in the beam may be integrated out. Alternately,some sort of synchronous detection system may be used, where thedetection is synchronized to the vibrating frequency.

Described herein are just some examples of the beneficial use ofnear-infrared or SWIR lasers for active remote sensing or hyper-spectralimaging. However, many other spectroscopy and identification procedurescan use the near-infrared or SWIR light consistent with this disclosureand are intended to be covered by the disclosure. As one example, thefiber-based super-continuum lasers may have a pulsed output with pulsedurations of approximately 0.5-2 nsec and pulse repetition rates ofseveral Megahertz. Therefore, the active remote sensing orhyper-spectral imaging applications may also be combined with LIDAR-typeapplications. Namely, the distance or time axis can be added to theinformation based on time-of-flight measurements. For this type ofinformation to be used, the detection system would also have to betime-gated to be able to measure the time difference between the pulsessent and the pulses received. By calculating the round-trip time for thesignal, the distance of the object may be judged. In another embodiment,GPS (global positioning system) information may be added, so the activeremote sensing or hyper-spectral imagery would also have a location tagon the data. Moreover, the active remote sensing or hyper-spectralimaging information could also be combined with two-dimensional orthree-dimensional images to provide a physical picture as well as achemical composition identification of the materials. These are justsome modifications of the active remote sensing or hyper-spectralimaging system described in this disclosure, but other techniques mayalso be added or combinations of these techniques may be added, andthese are also intended to be covered by this disclosure.

Although the present disclosure has been described in severalembodiments, a myriad of changes, variations, alterations,transformations, and modifications may be suggested to one skilled inthe art, and it is intended that the present disclosure encompass suchchanges, variations, alterations, transformations, and modifications asfalling within the spirit and scope of the appended claims.

While exemplary embodiments are described above, it is not intended thatthese embodiments describe all possible forms of the disclosure. Rather,the words used in the specification are words of description rather thanlimitation, and it is understood that various changes may be madewithout departing from the spirit and scope of the disclosure.Additionally, the features of various implementing embodiments may becombined to form further embodiments of the disclosure. While variousembodiments may have been described as providing advantages or beingpreferred over other embodiments with respect to one or more desiredcharacteristics, as one skilled in the art is aware, one or morecharacteristics may be compromised to achieve desired system attributes,which depend on the specific application and implementation. Theseattributes include, but are not limited to: cost, strength, durability,life cycle cost, marketability, appearance, packaging, size,serviceability, weight, manufacturability, ease of assembly, etc. Theembodiments described herein that are described as less desirable thanother embodiments or prior art implementations with respect to one ormore characteristics are not outside the scope of the disclosure and maybe desirable for particular applications.

What is claimed is:
 1. An optical tomography system, comprising: anarray of laser diodes to generate light having one or more opticalwavelengths that includes at least one near-infrared wavelength between600 nanometers and 1000 nanometers, at least one laser diode of thearray comprises one or more Bragg reflectors; at least one of the laserdiodes to pulse at a modulation frequency between 10 Megahertz and 1Gigahertz and to have a phase associated with the modulation frequency;wherein at least a portion of the light generated by the array isconfigured to penetrate tissue comprising skin; and a detection systemcomprising at least one photo-detector, a lens at an input to the atleast one photo-detector, and a processor to process digitized signalsreceived from the at least one photo-detector, the detection systemconfigured to (i) measure a phase shift of at least a portion of thelight from the array of laser diodes reflected from the tissue relativeto the portion of the light generated by the array to penetrate thetissue, (ii) measure time-of-flight of at least a portion of the lightfrom the array of laser diodes reflected from the tissue relative to theportion of the light generated by the array to penetrate the tissue,(iii) generate one or more images of the tissue based at least in parton an amplitude of at least a portion of the light from the array oflaser diodes reflected from the tissue, and (iv) detect oxy- ordeoxy-hemoglobin in the tissue.
 2. The optical tomography system ofclaim 1 wherein the detection system is configured to be synchronized tothe pulsing of the array of laser diodes, and wherein at least some ofthe laser diodes in the array of laser diodes operate at a wavelengthnear 940 nanometers.
 3. The optical tomography system of claim 1,wherein the detection system is configured to identify veins within thetissue.
 4. The optical tomography system of claim 1, wherein thedetection system is further configured to: receive light while the arrayof laser diodes is off and covert the received light into a firstsignal; receive light while at least a part of the array of laser diodesis on and convert the received light into a second signal, the receivedlight including at least a part of the light from the array of laserdiodes reflected from the tissue; and wherein the detection systemprocessor is configured to difference the first signal and the secondsignal.
 5. The optical tomography system of claim 1 further comprising:a second laser diode to be pulsed and to generate light having one ormore optical wavelengths, and wherein the second laser diode comprisesone or more Bragg reflectors; a beam splitter to: receive at least partof the light from the second laser diode, split the received light intoa sample arm and a reference arm, and direct at least a portion of thesample arm light to tissue comprising skin; at least one first detectorto receive at least a portion of the reference arm light and to generatea reference detector output; at least one second detector to receivefrom the tissue at least a portion of reflected sample arm light and togenerate a sample detector output; and wherein the detection systemprocessor is configured to measure another time-of-flight based at leastin part on the sample detector output.
 6. The optical tomography systemof claim 1, wherein the detection system processor is configured toimplement pattern matching and a threshold function to correlatedetected blood concentrations with a library of known concentrations todetermine overlap between the detected concentrations and knownconcentrations in the library.
 7. The optical tomography system of claim6, wherein the detection system processor is further configured toidentify differences in detected concentrations using one or more oflinear regression, partial least squares, and principal componentregression.
 8. A blood measurement system, comprising: an array of laserdiodes to generate light having one or more optical wavelengths thatincludes at least one near-infrared wavelength, the array of laserdiodes comprising one or more Bragg reflectors, wherein at least aportion of the light generated by the array is configured to penetratetissue comprising skin; at least one of the laser diodes to pulse at apulse repetition rate between one (1) kilohertz and about 100 megahertz;a detection system to non-invasively measure blood in veins based atleast in part on near-infrared diffuse reflection from the skin, thedetection system comprising at least one photo-detector and a lenssystem coupled to the at least one photo-detector, wherein the at leastone photo-detector is coupled to one or more analog-to-digitalconverters and a processor; the detection system configured to measureabsorption of hemoglobin in the near-infrared wavelength between 700nanometers and 1300 nanometers; and the detection system processorconfigured to: differentiate between a region in the skin having a veinand a region in the skin without distinct veins, and implement patternmatching and a threshold function to correlate detected bloodconcentrations with a library of known concentrations to determineoverlap between the detected blood concentrations and knownconcentrations in the library.
 9. The blood measurement system of claim8 coupled to a smart phone or tablet, wherein the smart phone or tabletfurther comprises a wireless receiver, a wireless transmitter, adisplay, a voice input module, and a speaker.
 10. The blood measurementsystem of claim 9, wherein the tissue comprises a hand of a user, andthe blood measurement system is used to measure veins in the hand. 11.The blood measurement system of claim 10, wherein the detection systemcomprises an infrared camera, wherein one or more optical filters andone or more adjustable lenses are configured to be coupled to thecamera, and wherein the detection system processor is configured toidentify differences in detected concentrations using one or more oflinear regression, partial least squares, and principal componentregression.
 12. The blood measurement system of claim 8, wherein thepattern matching comprises spectral fingerprinting, wherein the at leastone of the laser diodes emits a near-infrared wavelength near 940nanometers, and wherein the at least a portion of the light generated bythe array is configured to penetrate two or more spatial locations onthe skin.
 13. The blood measurement system of claim 8, comprising: asecond laser diode configured to be pulsed, and to generate light havingone or more optical wavelengths, wherein at least a portion of the oneor more optical wavelengths is a near-infrared wavelength between 700nanometers and 2500 nanometers, and wherein the second laser diodecomprises one or more Bragg reflectors; a beam splitter configured toreceive at least part of the light from the second laser diode, to splitthe light into a sample arm and a reference arm, and to direct at leasta portion of the sample arm light to tissue comprising skin; at leastone first detector configured to receive at least a portion of thereference arm light and configured to generate a reference detectoroutput; at least one second detector configured to receive from thetissue at least a portion of reflected sample arm light and configuredto generate a sample detector output; and wherein the detection systemprocessor is configured to measure a time-of-flight based at least inpart on the sample detector output.
 14. An optical identificationsystem, comprising: an array of laser diodes to generate light havingone or more optical wavelengths that includes at least one near-infraredwavelength between 600 nanometers and 1000 nanometers, at least onelaser diode of the array comprises one or more Bragg reflectors, whereinat least a portion of the light generated by the array of laser diodesis configured to be directed to tissue comprising skin; at least one ofthe laser diodes to pulse at a modulation frequency between 10 Megahertzand 1 Gigahertz and to have a phase associated with the modulationfrequency; and a detection system comprising at least onephoto-detector, a lens at an input to the at least one photo-detector,and a processor to process digitized signals received from the at leastone photo-detector, the detection system configured to (i) measure aphase shift of at least a portion of the light from the array of laserdiodes reflected from the tissue, (ii) measure time-of-flight of atleast a portion of the light from the array of laser diodes reflectedfrom the tissue, (iii) generate one or more images of the tissue basedat least in part on an amplitude of at least a portion of the light fromthe array of laser diodes reflected from the tissue, and (iv) generate adepth image from a temporal distribution of light relative to areference signal.
 15. The optical identification system of claim 14,wherein the detection system is configured to be synchronized to thepulsing of the array of laser diodes, and wherein at least some of thelaser diodes in the array of laser diodes are configured to operate at awavelength near 940 nanometers.
 16. The optical identification system ofclaim 14, wherein the detection system is configured to identify veinswithin the tissue.
 17. The optical identification system of claim 14,wherein the detection system is further configured to: receive lightwhile the array of laser diodes is off and covert the received lightinto a first signal; receive light while at least a part of the array oflaser diodes is on and convert the received light into a second signal,the received light including at least a part of the light from the arrayof laser diodes reflected from the tissue; and wherein the detectionsystem processor is configured to difference the first signal and thesecond signal.
 18. The optical identification system of claim 14,further comprising: a second laser diode configured to be pulsed, and togenerate light having one or more optical wavelengths, wherein at leasta portion of the one or more optical wavelengths is a near-infraredwavelength between 700 nanometers and 2500 nanometers, and wherein thesecond laser diode comprises one or more Bragg reflectors; a beamsplitter configured to: receive at least part of the light from thesecond laser diode, split the received light into a sample arm and areference arm, and direct at least a portion of the sample arm light totissue comprising skin; at least one first detector configured toreceive at least a portion of the reference arm light and to generate areference detector output; at least one second detector configured toreceive from the tissue at least a portion of reflected sample arm lightand to generate a sample detector output; and wherein the detectionsystem processor is configured to measure another time-of-flight basedat least in part on the sample detector output.
 19. The opticalidentification system of claim 14, wherein the detection systemprocessor is configured to implement pattern matching and a thresholdfunction to correlate detected blood concentrations with a library ofknown concentrations to determine overlap between the detected bloodconcentrations and known concentrations in the library.
 20. The opticalidentification system of claim 19, wherein the detection systemprocessor is configured to identify differences in detectedconcentrations using one or more of linear regression, partial leastsquares, and principal component regression.